q 2006 by Taylor & Francis Group, LLC - Developers

q 2006 by Taylor & Francis Group, LLC



q 2006 by Taylor & Francis Group, LLC 

Dedication 

To the loving memory of my dad, 

Mustafa Amiji, and to my mom, Shirin Amiji

Foreword 

We are at the threshold of a new era in cancer treatment and diagnosis, brought about by the 

convergence of two disciplines—materials engineering and life sciences—that 30 years ago might 

have been difficult to envision. The product of this curious marriage, nanobiotechnology, is yielding 

many surprises and fostering many hopes in the drug-development space. Nanoparticles, 

engineered to exquisite precision using polymers, metals, lipids, and carbon, have been combined 

with molecular targeting, molecular imaging, and therapeutic techniques to create a powerful set of 

tools in the fight against cancer. The unique properties of nanomaterials enable selective drug 

delivery to tumors, novel treatment methods, intraoperative imaging guides to surgery, highly 

sensitive imaging agents for early tumor detection, and real-time monitoring of response to 

treatment. 

Using nanotechnology, it may be possible to leap over many of the hurdles of cancer drug 

delivery that have confounded conventional drugs. These hurdles include hindered access to the 

central nervous system through the blood–brain barrier, sequestration by the reticulo-endothelial 

system, inability to penetrate the interior of solid tumors, and overcoming multi-drug resistance 

mechanisms. These obstacles can be mitigated by manipulating the size, surface charge, 

hydrophilicity, and attached targeting ligands of a therapeutic nanoparticle. 

The novelty of using nanotechnology for medical applications presents its own challenges, 

similar in many ways to the challenges faced by the introduction of the first synthetic protein-based 

drugs. In the early 1980s, protein-based drugs offered an entirely new approach to targeting and 

treating disease. They could be “engineered” to be highly specific and even offered the capability of 

separating the targeting and therapeutic functions of a drug, such as when monoclonal antibodies 

are conjugated to cytotoxic agents. But monoclonal antibodies and recombinant protein-based 

drugs necessitated a different approach to lead optimization, metabolism, and toxicity screening 

from that of small-molecule drugs. Factors such as stability, immunogenicity, and species 

specificity had to be considered and tested in ways that were not familiar to the developers of 

traditional drugs. 

Nanotechnology-based drugs will catalyze a reevaluation of optimization, metabolism, and 

toxicity-screening protocols. Interactions between novel materials and biological pathways are 

largely unknown but are widely suspected to depend heavily on physical and chemical 

characteristics such as particle size, particle size distribution, surface area, surface chemistry 

(including charge and hydrophobicity), shape, and aggregation state—features not usually 

scrutinized for traditional drugs. Standard criteria for physicochemical characterization will have to 

be established before safety testing can yield interpretable and reproducible results. 

Would nanotechnology-based drugs be successful? All drugs have to run an obstacle course 

through physiological and biochemical barriers. For monoclonal antibody drugs, only a minute 

portion of an intravenously administered drug reaches its target. However, drugs based on 

nanomaterials, including nanoparticles, have unique properties that enable them to specifically bind 

to and penetrate solid tumors. Size and surface chemistries can be manipulated to facilitate 

extravasation through tumor vasculature, or the therapeutic agent can be encapsulated in polymer 

micelles or liposomes to prevent degradation and increase circulation half-life to improve the odds 

of it reaching the target. 

This level of functional engineering has not been available to either small-molecule or proteinbased 

drugs. For these drugs, modification of one characteristic, such as solubility or charge, can 

have dramatic effects on other essential characteristics, such as potency or target specificity. 

Nanoparticles, however, introduce a much higher degree of modularity and offer at least four 

advantages over the antibody conjugates: (1) the delivery of a larger therapeutic payload per target 

recognition event, enhancing potency; (2) the ability to carry multiple targeting agents, enhancing 


selectivity; (3) the ability to carry multiple therapeutic agents, enabling targeted combination 

therapies; and (4) the ability to bypass physiological and biological barriers. 

We would all like to hasten the day when chemotherapy—the administration of non-specific, 

toxic anti-cancer agents—is relegated to medical history. Improvements in diagnostic screening 

and the development of drugs that target specific biological pathways have begun to turn the tide 

and contribute to a slow, yet consistent decline in death rates. While confirming that early diagnosis 

and targeted therapies are pointing us in the right direction, the progress is still incremental. 

Nanotechnology may be our best hope for overcoming many of the barriers faced by today’s drugs 

in the battle against cancer. The combined creative forces of engineering, chemistry, physics, and 

biology will beget new and hopefully transformative options to old, intractable problems. 


Piotr Grodzinski, Ph.D. 

National Cancer Institute

Preface 

With parallel breakthroughs occurring in molecular biology and nanoscience/technology, the newly 

recognized research thrust on “nanomedicine” is expected to have a revolutionary impact on the 

future of healthcare. To advance nanotechnology research for cancer prevention, diagnosis, and 

treatment, the United States National Cancer Institute (NCI) established the Alliance for 

Nanotechnology in Cancer in September 2004 and has pledged $144.3 million in the next five years 

(for details, visit http://nano.cancer.gov). Among the approaches for exploiting developments in 

nanotechnology for cancer molecular medicine, nanoparticles offer some unique advantages as 

sensing, delivery, and image enhancement agents. Several varieties of nanoparticles are available, 

including polymeric conjugates and nanoparticles, micelles, dendrimers, liposomes, and 

nanoassemblies. 

This book focuses specifically on nanoscientific and nanotechnological strategies that are 

effective and promising for imaging and treatment of cancer. Among the various approaches 

considered, nanotechnology offers the best promise for targeted delivery of drugs and genes to the 

tumor site and alleviation of the side effects of chemotherapeutic agents. Multifunctional 

nanosystems offer tremendous opportunity for combining more than one drug or using drug and 

imaging agents. The expertise of world-renowned academic and industrial researchers is brought 

together here to provide a comprehensive treatise on this subject. 

The book is composed of thirty-eight chapters divided into seven sections that address the 

specific nanoplatforms used for imaging and delivery of therapeutic molecules. Section 1 focuses 

on the rationale and fundamental understanding of targeting strategies, including pharmacokinetic 

considerations for delivery to tumors in vivo, multifunctional nanotherapeutics, boron neutron 

capture therapy, and the discussion on nanotechnology characterization for cancer therapy, as well 

as guidance from the U.S. Food and Drug Administration on approval of nanotechnology products. 

Section 2 focuses on polymeric conjugates used for tumor-targeted imaging and delivery, including 

special consideration on the use of imaging to evaluate therapeutic efficacy. In Section 3, polymeric 

nanoparticle systems are discussed with emphasis on biodegradable, long-circulating nanoparticles 

for passive and active targeting. Section 4 focuses on polymeric micellar assemblies, where 

sophisticated chemistry is applied for the development of novel nanosystems that can provide 

efficient delivery to tumors. Many of the micellar delivery systems are undergoing clinical trials in 

Japan and other countries across the globe. Dendritic nanostructures used for cancer imaging and 

therapy are discussed in Section 5. Section 6 focuses on the oldest nanotechnology for cancer 

therapy—liposome-based delivery systems—with emphasis on surface modification to enhance 

target efficiency and temperature-responsive liposomes. Lastly, Section 7 focuses on other lipid 

nanosystems used for targeted delivery of cancer therapy, including nanoemulsions that can cross 

biological barriers, solid-lipid nanoparticles, lipoprotein nanoparticles, and DQAsomes for 

mitochondria-specific delivery. 

Words cannot adequately express my admiration and gratitude to all of the contributing authors. 

Each chapter is written by a world-renowned authority on the subject, and I am deeply grateful for 

their willingness to participate in this project. I am also extremely grateful to Dr. Piotr Grodzinski 

for providing the Foreword. Drs. Fredika Robertson and Mauro Ferrari have done a superb job in 

laying the foundation by providing a chapter entitled “Introduction and Rationale for 

Nanotechnology in Cancer.” I am grateful to Professor Kinam Park at Purdue University, 

Professor Robert Langer at MIT, and Professor Vladimir Torchilin at Northeastern University, who 

have been my mentors and collaborators, as well as many other researchers from academia and 

industry. Special thanks are due to the postdoctoral associates and graduate students in my 

laboratory at Northeastern University who have been the “soldiers in the trenches” in our quest to 

use nanotechnology for the targeted delivery of drugs and genes to solid tumors. Lastly, I am deeply 


grateful to the wonderful people at Taylor & Francis-CRC Press, including Stephen Zollo, Patricia 

Roberson, and many others, who have made the concept of this book into reality. 

Any comments and constructive criticisms of the book can be sent to the editor at 

m.amiji@neu.edu. 


Editor 

Dr. Mansoor M. Amiji received his undergraduate degree in pharmacy from Northeastern 

University in 1988 and his PhD in pharmaceutics from Purdue University in 1992. His areas of 

specialization include polymeric biomaterials, advanced drug delivery systems, and nanomedical 

technologies. 

Dr. Amiji’s research interests include the synthesis of novel polymeric materials for medical and 

pharmaceutical applications; surface modification of cationic polymers by the complexationinterpenetration 

method to develop biocompatible materials; the preparation and characterization 

of polymeric membranes and microcapsules with controlled permeability properties for medical 

and pharmaceutical applications; target-specific drug and vaccine delivery systems for 

gastrointestinal tract infections; localized delivery of cytotoxic and anti-angiogenic drugs for 

solid tumors in novel biodegradable polymeric nanoparticles; intracellular delivery systems for 

drugs and genes using target-specific, long-circulating, biodegradable polymeric nanoparticles; and 

gold and iron-gold core-shell nanoparticles for biosensing, imaging, and delivery applications. His 

research has received sustained funding from the National Institutes of Health (NIH), the National 

Science Foundation (NSF), foundations, and local industries. 

Dr. Amiji is Professor and Associate Chair of the Pharmaceutical Sciences Department and Co- 

Director of the Northeastern University Nanomedicine Education and Research Consortium 

(NERC). The NERC oversees a doctoral training grant in nanomedicine science and technology 

that was co-funded by the NIH and NSF. He has two published books, Applied Physical Pharmacy 

and Polymeric Gene Delivery: Principles and Applications, along with numerous manuscript 

publications. He has also received a number of awards, including the 2003 Eurand Award for 

Innovative Oral Drug Delivery Research, Third Prize. 

Dr. Amiji has supervised the research efforts of over 50 postdoctoral associates, doctoral and 

master’s level graduate students, and undergraduate honors students over the last 13 years. His 

teaching responsibilities include the Doctor of Pharmacy (PharmD) program and graduate 

programs (MS and PhD) in pharmaceutical sciences, biotechnology, and nanomedicine. 


Contributors 

Hamidreza Montazeri Aliabadi 

Department of Pharmacy and Pharmaceutical 

Sciences 

University of Alberta 

Edmonton, Alberta, Canada 

Christine Allen 

Department of Pharmaceutical Sciences 

and 

Department of Chemical Engineering and 

Applied Chemistry 

University of Toronto 

Toronto, Ontario, Canada 

Ashootosh V. Ambade 

Department of Chemistry 

University of Massachusetts at Amherst 

Amherst, Massachusetts 

Mansoor M. Amiji 


Northeastern University 

Boston, Massachusetts 

Joseph M. Backer 

SibTech, Inc. 

Newington, Connecticut 

Marina V. Backer 

SibTech, Inc. 

Newington, Connecticut 

You Han Bae 

Department of Pharmaceutics and 

Pharmaceutical Chemistry 

University of Utah 

Salt Lake City, Utah 

Lajos P. Balogh 

Department of Radiation Medicine 

and 

Department of Cellular Stress and 

Cancer Biology 

Roswell Park Cancer Institute 

Buffalo, New York 


Rolf F. Barth 

Department of Pathology 

The Ohio State University 

Columbus, Ohio 

Reina Bendayan 




Tania Betancourt 

The University of Texas at Austin 

Austin, Texas 

D. Bhadra 


Dr. H.S. Gour University 

Sagar, India 

S. Bhadra 



Sagar, India 

Sangeeta N. Bhatia 

Harvard–MIT Division of Health Sciences 

and Technology 

Massachusetts Institute of Technology 

Cambridge, Massachusetts 

Sarathi V. Boddapati 




Lisa Brannon-Peppas 


Austin, Texas 

Mark Butler 

Department of Chemical Engineering 

and Applied Chemistry 



John C. Byrd 

Division of Hematology–Oncology 

NCI Comprehensive Cancer Center 


Columbus, Ohio

Robert B. Campbell 




Shing-Ming Cheng 




Gerard G. M. D’Souza 




Marina A. Dobrovolskaia 

Nanotechnology Characterization Laboratory 

NCI Frederick 

Frederick, Maryland 

Amber Doiron 


Austin, Texas 

P. K. Dubey 


Dr. H. S. Gold University 

Sagar, India 

Omid C. Farokhzad 

Department of Anesthesiology, Perioperative 

and Pain Medicine 

Brigham and Women’s Hospital and Harvard 

Medical School 


Mauro Ferrari 

The Brown Foundation Institute of 

Molecular Medicine 

M. D. Anderson Cancer Center 

and 

Alliance for Nanohealth 

The University of Texas 

Houston, Texas 

Alberto A. Gabizon 

Oncology Institute 

Shaare Zedek Medical Center and 

Hebrew University 

Jerusalem, Israel 


Jinming Gao 

Simmons Comprehensive Cancer Center 

University of Texas Southwestern 

Medical Center 

Dallas, Texas 

Hamidreza Ghandehari 


University of Maryland 

Baltimore, Maryland 

Todd J. Harris 





Mitsuru Hashida 

Department of Drug Delivery Research 

Graduate School of Pharmaceutical Sciences 

Kyoto University 

Sakyo-ku, Kyoto, Japan 

Yuriko Higuchi 





Kang Moo Huh 

Department of Polymer Science 

and Engineering 

Chungnam National University 

Yuseong-Gu, Daejeon, South Korea 

Edward F. Jackson 

Division of Diagnostic Imaging 

The University of Texas M. D. Anderson 

Cancer Center 


N. K. Jain 



Sagar, India 

Eun-Kee Jeong 

Department of Radiology 


Salt Lake City, Utah

Ji Hoon Jeong 

Department of Pharmaceutics and 




Sangyong Jon 

Department of Life Science 

Gwangju Institute of Science and Technology 

Gwangju, South Korea 

Shigeru Kawakami 





Mohamed K. Khan 

Department of Radiation Medicine 

and 

Department of Cellular Stress and 

Cancer Biology 

Roswell Park Cancer Institute 

Buffalo, New York 

Chalermchai Khemtong 




Dallas, Texas 

Sun Hwa Kim 

Department of Biological Sciences 

Korea Advanced Institute of Science 


Daejeon, South Korea 

Tae-il Kim 

School of Chemistry and Molecular 

Engineering 

Seoul National University 

Seoul, South Korea 

Sushma Kommareddy 




Vinod Labhasetwar 

Departments of Pharmaceutical Sciences and 

Biochemistry and Molecular Biology 

University of Nebraska Medical Center 

Omaha, Nebraska 


Andras G. Lacko 

Departments of Molecular Biology and 

Immunology 

University of North Texas Health Science 

Center 

Fort Worth, Texas 

Robert Langer 

Department of Chemical Engineering 



Afsaneh Lavasanifar 

Department of Pharmacy and Pharmaceutical 

Sciences 



Eun Seong Lee 

Amorepacific Corporation/R&D Center 

Giheung-gu Yongin-si Gyeonggi-do 

South Korea 

Helen Lee 




Robert J. Lee 


and 

NSF Nanoscale Science and 

Engineering Center 



Sang Cheon Lee 

Nanomaterials Application Division 

Korea Institute of Ceramic Engineering and 

Technology 

Guemcheon-gu, Seoul, South Korea 

Chun Li 

M. D. Anderson 

Cancer Center 



Yongqiang Li 



Toronto, Ontario, Canada

Bruce R. Line 

Department of Radiology, and Greenebaum 

Cancer Center 



Jubo Liu 




Zheng-Rong Lu 

Departments of Pharmaceutics and 




S. Mahor 


Dr. H. S. Gour University 

Sagar, India 

Walter J. McConathy 

Department of Internal Medicine 

University of North Texas Health Science 

Center 


Scott E. McNeil 


NCI Frederick 


T. J. Miller 

Center for Drug Evaluation and Research 

Food and Drug Administration 

Silver Spring, Maryland 

Amitava Mitra 

Department of Pharmaceutical Sciences and 

Center for Nanomedicine and 

Cellular Delivery 



S. M. Moghimi 

Molecular Targeting and Polymer 

Toxicology Group 

School of Pharmacy 

University of Brighton 

Brighton, United Kingdom 


Randall J. Mrsny 

Center for Drug Delivery/Biology 

Welsh School of Pharmacy 

Cardiff University 

Cardiff, United Kingdom 

R. S. R. Murthy 

Pharmacy Department 

The M.S. University of Baroda 

Vadodara, India 

Natarajan Muthusamy 




Anjan Nan 

Department of Pharmaceutical Sciences and 

Center for Nanomedicine and 

Cellular Delivery 



Maya Nair 

Departments of Molecular Biology and 

Immunology 

University of North Texas Health 

Science Center 


Norased Nasongkla 




Dallas, Texas 

David Needham 

Department of Mechanical Engineering 

and Materials Science 

Duke University 

Durham, North Carolina 

Tooru Ooya 

Department of Materials Science 

Japan Advanced Institute of Science 


Tatsunokuchi, Ishikawa, Japan

Jong-Sang Park 

Department of Chemistry and Molecular 

Engineering 

Seoul National University 

Seoul, South Korea 

Kinam Park 


Biomedical Engineering 

Purdue University 

West Lafayette, Indiana 

Tae Gwan Park 

Department of Biological Sciences 

Korea Advanced Institute of Science and 

Technology 

Daejeon, South Korea 

Anil K. Patri 


NCI Frederick 


Ana Ponce 

Department of Biomedical Engineering 

Duke University 

Durham, North Carolina 

Natalya Rapoport 

Department of Bioengineering 



Mike Andrew Rauth 




L. Harivardan Reddy 

Pharmacy Department 

The M.S. University of Baroda 

Vadodara, India 

Fredika M. Robertson 

College of Medicine and Comprehensive 

Cancer Center 




N. Sadrieh 

Center for Drug Evaluation and Research 

Food and Drug Administration 

Silver Spring, Maryland 

Sanjeeb K. Sahoo 


Nebraska Medical Center 

Omaha, Nebraska 

and 

Institute of Life Sciences 

Bhubaneswar, India 

Elamprakash N. Savariar 




Dinesh B. Shenoy 



Boston, Massachusetts and Novavax, Inc. 

Malvern, Pennsylvania 

Patrick Lim Soo 




Stephan T. Stern 

Nanotechnology Characterization 

Laboratory 

NCI Frederick 


S. (Thai) Thayumanavan 




Sandip B. Tiwari 




Werner Tjarks 

College of Pharmacy 


Columbus, Ohio

Vladimir P. Torchilin 




Anagha Vaidya 





Geoffrey von Maltzahn 





S. P. Vyas 


Dr. H. S. Gour University 

Sagar, India 

Sidney Wallace 

Division of Diagnostic Imaging 

The University of Texas M. D. Anderson 

Cancer Center 


Yanli Wang 





Volkmar Weissig 





Ho Lun Wong 




Gong Wu 




Xiao Yu Wu 




Xiao-Bing Xiong 

Faculty of Pharmacy and Pharmaceutical 

Sciences 



Weilian Yang 




Furong Ye 





Guodong Zhang 

M. D. Anderson 

Cancer Center 



Xiaobin B. Zhao 




.

Table of Contents 

Section 1 

Nanotechnology and Cancer. .................................................1 

Chapter 1 Introduction and Rationale for Nanotechnology 

in Cancer Therapy . ..............................................3 

Fredika M. Robertson and Mauro Ferrari 

Chapter 2 Passive Targeting of Solid Tumors: Pathophysiological 

Principles and Physicochemical Aspects of 

Delivery Systems . . .............................................11 

S. M. Moghimi 

Chapter 3 Active Targeting Strategies in Cancer with a Focus on 

Potential Nanotechnology Applications . ..............................19 


Chapter 4 Pharmacokinetics of Nanocarrier-Mediated 

Drug and Gene Delivery . . . ......................................43 

Yuriko Higuchi, Shigeru Kawakami, and Mitsuru Hashida 

Chapter 5 Multifunctional Nanoparticles for Cancer Therapy. . . . ...................59 

Todd J. Harris, Geoffrey von Maltzahn, and Sangeeta N. Bhatia 

Chapter 6 Neutron Capture Therapy of Cancer: Nanoparticles 

and High Molecular Weight Boron Delivery Agents. . . ...................77 

Gong Wu, Rolf F. Barth, Weilian Yang, Robert J. Lee, Werner Tjarks, Marina V. Backer, and 

Joseph M. Backer 

Chapter 7 Preclinical Characterization of Engineered Nanoparticles 

Intended for Cancer Therapeutics . .................................105 

Anil K. Patri, Marina A. Dobrovolskaia, Stephan T. Stern, and Scott E. McNeil 

Chapter 8 Nanotechnology: Regulatory Perspective for Drug 

Development in Cancer Therapeutics . . .............................139 

N. Sadrieh and T. J. Miller 

Section 2 

Polymer Conjugates. . ....................................................157 

Chapter 9 Polymeric Conjugates for Angiogenesis Targeted Tumor 

Imaging and Therapy . . . . . . .....................................159 

Amitava Mitra, Anjan Nan, Bruce R. Line, and Hamidreza Ghandehari 


Chapter 10 Poly(L-Glutamic Acid): Efficient Carrier of Cancer Therapeutics and 

Diagnostics . . . . . ............................................185 

Guodong Zhang, Edward F. Jackson, Sidney Wallace, and Chun Li 

Chapter 11 Noninvasive Visualization of In Vivo Drug Delivery 

of Paramagnetic Polymer Conjugates with MRI. . .....................201 

Zheng-Rong Lu, Yanli Wang, Furong Ye, Anagha Vaidya, and Eun-Kee Jeong 

Section 3 

Polymeric Nanoparticles . . . . . . ............................................213 

Chapter 12 Polymeric Nanoparticles for Tumor-Targeted Drug 

Delivery. ...................................................215 

Tania Betancourt, Amber Doiron, and Lisa Brannon-Peppas 

Chapter 13 Long-Circulating Polymeric Nanoparticles for Drug and 

Gene Delivery to Tumors . . . ....................................231 

Sushma Kommareddy, Dinesh B. Shenoy, and Mansoor M. Amiji 

Chapter 14 Biodegradable PLGA/PLA Nanoparticles for Anti-cancer 

Therapy . ...................................................243 

Sanjeeb K. Sahoo and Vinod Labhasetwar 

Chapter 15 Poly(Alkyl Cyanoacrylate) Nanoparticles for Delivery 

of Anti-Cancer Drugs . . ........................................251 

R. S. R. Murthy and L. Harivardhan Reddy 

Chapter 16 Aptamers and Cancer Nanotechnology .............................289 

Omid C. Farokhzad, Sangyong Jon, and Robert Langer 

Section 4 

Polymeric Micelles . . . ...................................................315 

Chapter 17 Polymeric Micelles for Formulation of Anti-Cancer Drugs. . . . . ..........317 

Helen Lee, Patrick Lim Soo, Jubo Liu, Mark Butler, and Christine Allen 

Chapter 18 PEO-Modified Poly(L-Amino Acid) Micelles for 

Drug Delivery . . . ............................................357 

Xiao-Bing Xiong, Hamidreza Montazeri Aliabadi, and Afsaneh Lavasanifar 

Chapter 19 Hydrotropic Polymer Micelles for Cancer Therapeutics .................385 

Sang Cheon Lee, Kang Moo Huh, Tooru Ooya, and Kinam Park 

Chapter 20 Tumor-Targeted Delivery of Sparingly-Soluble Anti-Cancer 

Drugs with Polymeric Lipid-Core Immunomicelles . . . .................409 



Chapter 21 Combined Cancer Therapy by Micellar-Encapsulated 

Drugs and Ultrasound. . . . . .....................................421 


Chapter 22 Polymeric Micelles Targeting Tumor pH. . . . . . ......................443 

Eun Seong Lee and You Han Bae 

Chapter 23 cRGD-Encoded, MRI-Visible Polymeric Micelles 

for Tumor-Targeted Drug Delivery . . . .............................465 

Jinming Gao, Norased Nasongkla, and Chalermchai Khemtong 

Chapter 24 Targeted Antisense Oligonucleotide Micellar 

Delivery Systems . ............................................477 

Ji Hoon Jeong, Sun Hwa Kim, and Tae Gwan Park 

Section 5 

Dendritic Nanocarriers. ...................................................487 

Chapter 25 Dendrimers as Drug and Gene Delivery Systems ......................489 

Tae-il Kim and Jong-Sang Park 

Chapter 26 Dendritic Nanostructures for Cancer Therapy . . ......................509 

Ashootosh V. Ambade, Elamprakash N. Savariar, and S. (Thai) Thayumanavan 

Chapter 27 PEGylated Dendritic Nanoparticulate Carriers of 

Anti-Cancer Drugs ............................................523 

D. Bhadra, S. Bhadra, and N. K. Jain 

Chapter 28 Dendrimer Nanocomposites for Cancer Therapy ......................551 

Lajos P. Balogh and Mohamed K. Khan 

Section 6 

Liposomes . . . . ........................................................593 

Chapter 29 Applications of Liposomal Drug Delivery Systems 

to Cancer Therapy ............................................595 


Chapter 30 Positively-Charged Liposomes for Targeting 

Tumor Vasculature. ...........................................613 


Chapter 31 Cell Penetrating Peptide (CPP)–Modified Liposomal 

Nanocarriers for Intracellular Drug and Gene Delivery . . . ..............629 



Chapter 32 RGD-Modified Liposomes for Tumor Targeting . .....................643 

P. K. Dubey, S. Mahor, and S. P. Vyas 

Chapter 33 Folate Receptor-Targeted Liposomes for 

Cancer Therapy . . ............................................663 

Xiaobin B. Zhao, Natarajan Muthusamy, John C. Byrd, and Robert J. Lee 

Chapter 34 Nanoscale Drug Delivery Vehicles for Solid Tumors: 

A New Paradigm for Localized Drug Delivery 

Using Temperature Sensitive Liposomes . . . .........................677 

David Needham and Ana Ponce 

Section 7 

Other Lipid Nanostructures . . . . ............................................721 

Chapter 35 Nanoemulsion Formulations for Tumor-Targeted Delivery . . . . . ..........723 

Sandip B. Tiwari and Mansoor M. Amiji 

Chapter 36 Solid Lipid Nanoparticles for Antitumor Drug Delivery .................741 

Ho Lun Wong, Yongqiang Li, Reina Bendayan, Mike Andrew Rauth, and Xiao Yu Wu 

Chapter 37 Lipoprotein Nanoparticles as Delivery Vehicles 

for Anti-Cancer Agents ........................................777 

Andras G. Lacko, Maya Nair, and Walter J. McConathy 

Chapter 38 DQAsomes as Mitochondria-Targeted Nanocarriers 

for Anti-Cancer Drugs . ........................................787 

Shing-Ming Cheng, Sarathi V. Boddapati, Gerard G. M. D’Souza, and Volkmar Weissig 


1 

CONTENT 

Introduction and Rationale 

for Nanotechnology 

in Cancer Therapy 

Fredika M. Robertson and Mauro Ferrari 

References ........................................................................................................................................ 8 

Although there have been significant advances in defining the fundamentals of cancer biology over 

the past 25 years, this has not translated into similar clinical advances in cancer therapeutics. One 

area that holds great promise for making such advances is the area defined as cancer nanotechnology, 

which involves the intersection of a variety of disciplines, including engineering, materials 

science, chemistry, and physics with cancer biology. This multidisciplinary convergence has 

resulted in the creation of devices and/or materials that are themselves or have essential components 

in the 1–1000-nm range for at least one dimension and holds the possibility of rapidly 

advancing the state of cancer therapeutics and tumor imaging. 

This newly developing area of “nanohealth” may ultimately allow detection of human tumors at 

the very earliest stages, regardless of the location of the primary tumor and/or metastases, and may 

provide approaches to more effectively destroy tumors as well as their associated vascular supplies 

with fewer adverse side effects. The chapters within this book describe the most recent cutting-edge 

approaches in nanotechnology that are focused on overcoming the multiple barriers that have, in the 

past, blocked the successful treatment and ultimate eradication of human cancers. 

The majority of nanotechnology-based devices useful for cancer therapeutics have been defined 

as nanovectors, which are injectable nanoscale delivery systems. 1,2 Nanovectors offer the promise 

of providing breakthrough solutions to the problems of optimizing efficacy of therapeutic agents 

while simultaneously diminishing the deleterious side-effects that commonly accompany the use of 

both single chemotherapeutic agents as well as multimodality therapeutic regimens. 

In general, nanovectors are comprised of at least three constituents, which include a core 

material, a therapeutic and/or imaging “payload,” and a biological surface modification, which 

aids in both appropriate biodistribution and selective localization of the nanovector and its cytotoxic 

and/or imaging agent. Although the first type of molecules used to enhance the selective 

localization and delivery of nanovectors were antibodies, more sophisticated recognition systems 

have been devised as a result of our expanding knowledge base in cancer biology. 

For example, while the biological and molecular characteristics of human tumors of different 

origins continue to be defined and exploited for cancer therapeutics and tumor imaging, the significance 

of blood vessels that develop around actively growing tumors as potential therapeutic 


3

4 

Nanotechnology for Cancer Therapy 

targets has only been widely recognized and found clinical utility within the past decade. 3,4 The 

process of tumor-associated angiogenesis is now known to be an essential component of tumor 

expansion and metastasis. This realization and the accompanied potential for development of 

successful cancer therapeutics, as well as for tumor imaging strategies based on tumor-associated 

vasculature, has opened new doors for the application of nanotechnology in cancer therapeutics. 

Additionally, the molecules that drive the process of tumor angiogenesis may provide a means to 

gauge the timing and extent of individual patient responses to cancer treatment, which can also be 

monitored using nanovector approaches. The importance of tumor neovasculature in cancer nanotechnology 

is highlighted throughout the chapters in Nanotechnology for Cancer Therapy. 

One of the most important characteristics of nanovectors is their ability to be functionalized to 

overcome barriers that block access of agents used for treatment of tumors and for imaging of 

tumors and their associated vasculature. These biological barriers are numerous and complex. One 

such barrier is the blood–brain barrier, which prevents access to brain malignancies, compounding 

the difficulties in their successful treatment. One example of the potential utility of nanovectors in 

overcoming this biobarrier, which is critical to treatment of malignant brain tumors, is the use of 

nanoparticles in combination with boron neutron capture therapy (BCNT). The current approaches 

used for BCNT combining nanoparticles and the potential for this nanotechnology-based approach 

to treatment of brain tumors is described in detail in a chapter in Nanotechnology for Cancer 

Therapy. Additional biobarriers that must be overcome include epithelial–endothelial cell barriers, 

the barriers presented by the markedly tortuous structures that are characteristic of angiogenic 

vasculature associated with tumors, as well as the barrier set up by the rapid uptake of nanovectors 

by resident macrophages within the reticuloendothelial system (RES) which may prevent nanovectors 

from reaching their targeted location. 

To achieve breakthrough advances in cancer therapeutics, there are two related and essential 

components which must be addressed. The first issue in successful use of nanovectors is recognition 

of the tumor and the second is the ability of the nanovector to reach the site of the tumor and 

associated blood vessels. The goal is to preferentially achieve high concentrations of a specific 

chemotherapeutic agent, a tumor imaging agent, and/or gene therapies at the site(s) of tumors and 

associated vasculature. In addition, nanovectors must be able to deliver an active agent to achieve 

effective anti-tumor treatment, or tumor imaging, which is essential for tumor diagnosis and for 

monitoring the extent and timing of an individual patient’s response to anti-tumor therapy. 

The first nanotechnology-based approach to be used as a means of delivering cancer 

chemotherapy was liposomes, which is a type of nanovector made of lipids surrounding a water 

core. Liposomes are the simplest form of nanovector and their utility is based on the significant 

difference in endothelial structures—defined as fenestrations—between normal vasculature and 

tumor-associated vessels. The increase in fenestrations in tumor neovasculature allows the preferential 

concentration of liposome-encapsulated anti-tumor agent in close proximity to the local 

tumor site, a phenomenon defined as enhanced penetration and retention (EPR), which is 

considered to be a characteristic of passive targeting of tumors. Both the passive targeting of 

solid tumors as well as approaches for active targeting of nanovectors such as liposomes are 

described in the following chapters. 

The first studies to demonstrate the proof of principle that liposomes could be used as a 

nanovector for effective delivery of an active cancer therapeutic was based on the encapsulation 

of the anthracycline antibiotic doxorubicin, which is an effective and commonly used first line 

chemotherapeutic agent for both leukemia and solid tumors. Although use of doxorubicin is 

associated with dose limiting cardiotoxicity, when anionic liposomes were used to encapsulate 

doxorubicin, preclinical studies found that the anti-tumor action of doxorubicin was significantly 

enhanced and the cardiotoxicity was significantly diminished. 5,6 Liposome encapsulated doxorubicin 

has now been clearly demonstrated safe and efficacious clinically in such malignancies as 

breast and ovarian cancer, 7–9 allowing for greater amounts of this agent to be administered than the 

recommended lifetime cumulative dose of free doxorubicin, which was 450–550 mg/m 2 . 


Introduction and Rationale for Nanotechnology in Cancer Therapy 5 

Importantly, liposomes have been shown to be an effective means of delivering a diverse group of 

anti-tumor agents such as doxorubicin, as well as the poorly soluble drug paclitaxel. 10 

As evidenced by the chapters in this volume, interest in improving the use of liposomes for 

enhancing the selective localization and delivery of cancer therapeutics to tumors and tumorassociated 

blood vessels remains an active area of investigation. Because liposomes can be 

modified with respect to composition, charge, and external coating, liposomes remain a versatile 

nanotechnology platform, historically serving as the prototypic nanovector. 

An approach that was first used with liposome nanovectors that has found utility with other 

types of nanovectors is “PEGylation,” a modification of liposome surface characteristics using 

poly(ethylene glycol) (PEG), resulting in what has been defined as “sterically stabilized liposomes.” 

This PEG modification of liposomes provides protection against uptake by resident 

macrophages within the RES biobarrier, thus increasing the circulation time of liposome-encapsulated 

anti-tumor agent, resulting in significantly increased therapeutic efficacy. 11 

This approach to overcoming rapid uptake and destruction by resident macrophages within the 

RES by PEGylation has been used alone or in tandem with other modifications of liposomes to aid 

in avoiding or overcoming biobarriers and to more selectively localize nanovectors. For example, 

based on the recognition that a wide variety of solid tumors as well as hematologic malignancies 

express amplified levels of the folate receptor on their cell surface, liposome nanovectors have been 

developed to target folate receptors. 12,13 Using the strategy of biomarker-based targeting, folate 

receptor targeted liposomes have been shown to not only overcome multidrug resistance associated 

with P-glycoprotein that often occurs in human tumors, but also have proven to be antiangiogenic. 13 

Another means of more selectively localizing liposome-encapsulated anti-tumor agents to tumorassociated 

vasculature is to add a cationic charge, 10,14 which has been shown to double the 

accumulation of anti-tumor drug encapsulated into liposomes selectively within vessels 

surrounding tumors. Other approaches to biomolecular targeting have included the use of specific 

proteins such as vasoactive intestinal peptide (VIP) 15 to target gastric tumors as well as signature 

RGD amino acid sequences expressed on the surface of endothelial cells that are associated with 

expression of a vb 3 integrin on angiogenic endothelium. 16 

In addition to serving as cancer drug delivery systems, liposomes have been shown to be an 

effective means for delivery of other agents such as genes and antisense oligonucleotides, and 

would allow access of such entities as small interfering RNA. 13,17 As an example, peptides such as 

cell penetrating peptides (CPPs) have been used to modify liposomes, allowing them to be used as a 

means for intracellular drug delivery and delivery of proteins such as antibodies and genes, as well 

as providing a window for cellular imaging. 17 

In addition to liposomes as an example of the prototype of nanovectors, there are many others 

now available in the nanotechnology toolbox that are being investigated for use in cancer therapeutics 

and tumor imaging. Polymer-based nanovectors are a specific area of interest in cancer 

therapeutics and a number of polymer-based nanovector systems are described in the following 

chapters. For the purposes of this volume, polymer-based nanovectors have been subdivided into 

several categories, including polymer conjugates, polymeric nanoparticles, and polymeric micelles. 

Polymeric conjugates that have been investigated for use as nanovector systems for 

delivery of anti-tumor agents include the water soluble biocompatible polymer N-(2hydroxypropyl)methacrylamide 

(HPMA). This polymeric conjugate has been targeted based on 

overexpression of hyaluronan (HA) receptors that are present on cancer cells. HPMA–HA polymeric 

conjugate drug delivery nanovector systems have been developed to carry a doxorubicin “payload” 

and have been shown to have selectivity for endocytosis of the targeted polymeric nanovector to 

breast, ovarian, and colon tumor cells, compared to an HPMA polymer with doxorubicin but lacking 

the HA targeting conjugate. 18,19 Copolymer–peptide conjugates have also been developed as nanovectors 

using HPMA in combination with RGD targeting peptides, 20 which molecularly target 

radiotherapeutic agents to tumor-associated vasculature, resulting in both antiangiogenic as well 

as anti-tumor activity. In addition to serving as nanovectors for delivery of chemotherapy and 


6 


radiotherapy, copolymer nanovectors show promise as agents to deliver biodegradable blood-pool 

contrast agents for use with tumor imaging methodologies. Other copolymers that have been used 

with gadolinium chelates for improvements in MRI based contrast agents include poly(L-glutamic 

acid) 21 as well as (Gd-DTPA)-cystine copolymers (GDCP) and (Gd-DTPA)-cystine diethyl ester 

copolymers (GDCEP). 22,23 As with liposome nanovectors, polymer conjugates can be targeted to a 

specific site, can carry a multitude of agents useful for cancer chemotherapeutics, radiotherapy, or 

tumor imaging, and their surfaces can be modified using different biological surface modifiers to 

enhance selectivity of localization of the nanovector. 

Studies focusing on polymeric nanoparticles as nanovectors have included the use of poly(lacticco-glycolic 

acid) (PLGA) and polylactide (PLA) that can be selectively targeted against such surface 

receptors as the transferrin receptor on breast tumor cells, as well as can be optimized to avoid 

biobarriers such as the RES. In addition, polymeric nanoparticles can be altered to enhance the 

intracellular drug concentration once delivered. Polymeric nanoparticles are also mutable for 

enhancement of retention at the local tumor site and to combine anti-tumor agents with antiangiogenesis 

strategies. 24–26 In addition to delivery of anti-tumor agents that are antiangiogenic, 

nanovectors such as PLGA and PLA polymers as well as PEGylated gelatin nanoparticles have 

been developed. In the case of PEGylated gelatin nanovectors, these have been improved by thiolation 

to optimize their time in circulation and to enhance delivery of payloads such as non-viral DNA 

for gene therapy. 25,27–29 

Relatively new polymeric nanovectors recently shown to have potential utility in cancer nanotechnology 

are bioconjugates composed of PLGA polymers and nucleic acid ligands defined as 

aptamers. 30 Because these nanovectors can be PEGylated and further modified with selectively 

targeted approaches directed against such proteins as prostate specific antigen (PSA) which are 

present in abundance on the surface of prostatic tumor cells, these nanoparticle–aptamer bioconjugates 

represent a new class of polymeric nanoparticles with promise as multifunctional 

nanovectors. Polymeric aptamer bioconjugates are currently being evaluated for their optimization 

for use in a large number of systems that may lead to in vivo studies in the near future. 31 

Another type of nanovector discussed in Nanotechnology for Cancer Therapy includes polymeric 

micelles, which can be defined as self-assembling nanosized colloidal particles with a 

hydrophobic core and hydrophilic shell. 32 Polymeric micelles have been commonly used in the 

field of drug delivery for over two decades. The advantages of these nanovectors include the ability 

to encapsulate agents with poor aqueous solubility profiles such as has been observed with the antitumor 

agent paclitaxel 33 as well as with the photodynamic therapy agent, meso-tetratphenylporphine 

(TPP). 34 Other advantages associated with the use of polymeric micelles include the ability to 

control both drug targeting using immunotargeting 34 and rates of drug release. 32,35 The efficacy of 

agents encapsulated within polymeric micelle nanovectors occurs via the passive targeting method 

defined as enhanced penetration and retention (EPR), which results in diminished adverse side 

effects usually observed in normal tissues by administration of cytotoxic cancer 

chemotherapeutic agents. 

In addition to the utility of polymeric micelles for delivery of anti-cancer agents, 32–35 these 

nanovectors are also described as an effective means for the combined delivery of anti-tumor agents 

as well as simultaneous tumor imaging using focused ultrasound imaging. This makes it possible to 

non-invasively penetrate tumors that are typically difficult to access, such as with ovarian cancer. 36 

Polymeric micelles have also found usefulness as multifunctional nanovectors that are both visible 

to MRI imaging modalities using RGD-based targeting of the tumor-associated vasculature 

combined with drug delivery. Other approaches using polymeric micelles have focused on targeting 

the acidic extracellular pH of the tumor microenvironment. 37 Polymeric micelles have also been 

evaluated for delivery of antisense oligonucleotides using folate receptor targeting strategies. 38 

Although polymeric micelle nanovectors have been used as strategies for drug delivery, it is clear 

that they continue to have promise in cancer nanotechnology. 



Another class of nanovectors described in the following chapters is dendrimers, which are self 

assembling synthetic polymers possessing tunable nanoscale dimensions that assume conformations 

that are dependent upon the microenvironment. The development of polymeric micelles 

preceded the development of dendrimers. However, due to the characteristics of dendrimers as 

“perfectly branched monodispersed macromolecules,” 39 this class of nanovectors may be suitable 

as exquisitely sensitive carriers for defined release of anti-tumor agents, as well as very useful for 

targeted delivery of such molecules as contrast agents for tumor imaging. 

Dendritic nanocarriers have been evaluated for their capability to serve as nanovectors for 

delivery of drugs and for plasmid DNA, suggesting that dendrimer nanovectors may be useful 

for gene therapy. 40 Like liposomes, dendrimers can also be modified by PEGylation 41 to increase 

their bioavailability and half-life in the circulation. Dendrimers have been shown to be suitable for 

enhancing the solubility of drugs, and to have suitable profiles for controlled delivery as well as 

selective delivery of drugs such as flurbiprofen that are active within a local site of inflammation. 42 

Although dendrimers that have been evaluated for potential utility in cancer therapeutics have 

primarily been composed of polyamidoamine (PAMAM), there are also gold/dendrimer nanocomposites 

43 and silver/dendrimer nanocomposite materials 44 that have been developed and are 

currently being evaluated for their comparative chemical and biological characteristics with 

potential utility as biomarkers, as well as for use in cancer therapeutics, tumor imaging, and 

radiotherapy. 

The varied nature of nanovectors covered in Nanotechnology for Cancer Therapy is illustrated 

by such novel classes of nanocarriers as DQAsomes, which are mitochondrial-targeted nanovectors 

capable of delivery of anti-cancer drugs as well as carriers of gene therapy directed against this 

organelle. 44,45 The development of this nanovector delivery approach is based on the relatively 

recent realization that a number of genetic disorders associated with defects in oxidative metabolism 

are associated with genetic defects in mitochondrial DNA, with a significant lack of available 

therapies. While the majority of nanovectors target tumors and/or tumor-associated vasculature, the 

DQAsomes are composed of derivatives of the self-assembling mitochondriotropic bola-amphiphile 

dequalinium chloride, which forms cationic vesicles which can bind and transport DNA to 

mitochondria in living mammalian cells. 44 

Another approach using selective targeting of more classically used characteristics of cancer 

cells includes the development of lipoprotein nanoparticles as delivery vehicles for anti-cancer 

therapy. 46 These high density lipoprotein (rHDL) nanoparticles are composed of phosphatidylcholine, 

apolipoprotein A-1, cholesterol, and cholesteryl esters with encapsulated paclitaxel. 

They were shown to be useful as a nanovector for anti-tumor drug delivery based on the ability 

of cancer cells to selectively acquire HDL, while the use of these nanovectors decreased the side 

effects normally observed with this type of anti-tumor therapy. 

Given the wide variety of nanovectors and nanoplatforms, Nanotechnology for Cancer Therapy 

cannot possibly cover each nanoscale application that has potential for use in anti-tumor therapy 

and tumor imaging. One nanoplatform that should be mentioned is the so-called “quantum dots” 

(Q-dots) that have recently received wide attention. Although cadmium and selenium crystalline 

nanovectors were first used over two decades ago, Q-dots have recently received attention for their 

utility as tunable multicolor nanoparticles that can be used for in vivo tumor imaging in animal 

models of cancer as well as for multiplex detection of mRNAs and proteins. Although the current 

composition of the Q-dots nanovectors prohibits their clinical use for cancer therapy, there is 

significant promise for the potential for development of biocompatible nanovectors which are 

tunable to wavelengths useful for tumor imaging. They may provide the ability to selectively 

target tumors as well as serve as multifunctional nanovector devices for cancer therapy, 

imaging, diagnostics, and monitoring of response to anti-tumor therapy. 

Although the development of nanoplatforms for cancer therapeutics is a critical component to 

the successful use of nanovectors for anti-tumor therapy and tumor imaging, there are other necessary 

components for nanoplatforms to gain clinical acceptance. One such parameter is the necessary 


8 

infrastructure to characterize nanoplatforms and to ensure that the rapid translation of discoveries in 

this area proceeds from preclinical to clinical use. The recognition by the National Cancer Institute 

(NCI) that nanodevices and nanomaterials will be critical in making significant advances in 

imaging, diagnosis, and treatment of human tumors has lead to the development of The Nanotechnology 

Characterization Laboratory (NCL). This laboratory provides critical support to the NCI’s 

Alliance in Nanotechnology for Cancer. The role and function of the services provided by NCL is 

supported by the NCI, and works in concert with the National Institute of Standards and Technology 

(NIST) and the U.S. Food and Drug Administration (FDA). The NCL provides a valuable 

service to those investigators working in the area of cancer nanotechnology and performs preclinical 

efficacy and toxicity testing of nanoscale devices and materials. The goal of the NCL is to 

support the rapid acceleration of preclinical development of nanovectors to support IND application 

and subsequent clinical trials. 

The continued development and ultimate success of nanotechnology-based strategies for cancer 

therapeutics, imaging, and tracking patient therapeutic response will be dependent upon a number 

of parameters. A multidisciplinary approach to achieving success in cancer nanotechnology is 

absolutely necessary. While remaining grounded in our basic knowledge of the known targets 

that are critical for localization of multifunctional nanovectors to the tumor and surrounding 

tissue, researchers must be aware of new targets that have been identified as being critically 

involved in the response of tumors and tumor vasclulature to anti-tumor agents, and to recognize 

and take full advantage of known targets such as surface receptors that are amplified on tumor cells 

as well as recognize and incorporate newer targets such as mitochondrial DNA and HDL uptake. 

For example, while researchers continue to identify novel means to functionalize nanovectors to 

eliminate biobarriers that historically have limited access to tumors and their associated angiogenic 

vasculature and stroma, we will still need to produce and test multifunctional nanovectors capable 

of selective localization and delivery of agents that function both as cytotoxic agents as well as 

imaging agents. 

Because the size of nanodevices and nanomaterials is similar to that of naturally occurring 

components of cells, including surface receptors, organelles such as mitochondria, lipoproteins, and 

other, as yet unidentified molecules such as siRNA, microRNA, genes, and unknown proteins, 

nanodevices and nanomaterials can easily interface with biological molecules. The mutability of 

nanotechnology platforms has allowed their rapid introduction into the field of basic cancer biology 

and the specialty of oncology. The approaches described in the chapters within Nanotechnology for 

Cancer Therapy range from liposomes, which are among the first and the simplest nanovectors to 

be used for delivery of cytotoxic drugs, to advanced nanoscale devices for use as imaging contrast 

agents in conjunction with ultrasound and magnetic resonance imaging (MRI). Each of the nanotechnology 

approaches described in this book will be absolutely necessary to achieve significant 

breakthroughs in cancer therapeutics. It is truly an exciting time to be an integral part of this field as 

these discoveries in cancer nanotechnology are made and put into practice with the goal to use 

nanovectors and nanotechnology platforms to significantly impact cancer morbidity and mortality 

in the next decade. 

REFERENCES 


1. Ferrari, M., Nanovector therapeutics, Curr. Opin. Chem. Biol., 9, 343–346, 2005. 

2. Ferrari, M., Cancer nanotechnology: Opportunities and challenges, Nat. Rev. Cancer, 5(3), 161–171, 

2005. 

3. Folkman, J., Fundamental concepts of the angiogenic process, Curr. Mol. Med., 3(7), 643–651, 2003. 

4. Ferrara, N. and Kerbel, R. S., Angiogenesis as a therapeutic target, Nature, 438, 967–974, 2005. 

5. Treat, J., Greenspan, A., Forst, D., Sanchez, J. A., Ferrans, V. J., Potkul, L. A., Woolley, P. V., and 

Rahman, A., Antitumor activity of liposome-encapsulated doxorubicin in advanced breast cancer: 

Phase II study, J. Natl. Cancer Inst., 82(21), 1706–1710, 1990. 



6. Forssen, E. A. and Tokes, Z. A., Improved therapeutic benefits of doxorubicin by entrapment in 

anionic liposomes, Cancer Res., 43(2), 546–550, 1983. 

7. Straubinger, R. M., Lopez, N. G., Debs, R. J., Hong, K., and Papahadjopoulos, D., Liposome-based 

therapy of human ovarian cancer: Parameters determining potency of negatively charged and 

antibody-targeted liposomes, Cancer Res., 48(18), 5237–5245, 1988. 

8. Robert, N. J., Vogel, C. L., Henderson, I. C., Sparano, J. A., Moore, M. R., Silverman, P., Overmoyer, 

B. A. et al., The role of the liposomal anthracyclines and other systemic therapies in the management 

of advanced breast cancer, Semin. Oncol., 31(6 Suppl 13), 106–146, 2004. 

9. Ewer, M. S., Martin, F. J., and Henderson, Use of anionic liposomes for the reduction of chronic 

doxorubicin-induced cardiotoxicity, Proc. Natl. Acad. Sci. USA, 78(3), 1873–1877, 1981. 

10. Campbell, R. B., Balasubramanian, S. V., and Straubinger, R. M., Influence of cationic lipids on the 

stability and membrane properties of paclitaxel-containing liposomes, J. Pharm. Sci., 90(8), 

1091–1105, 2001. 

11. Gabizon, A., Shmeeda, H., Horowitz, A. T., and Zalipsky, S., Tumor cell targeting of liposomeentrapped 

drugs with phospholipid-anchored folic acid-PEG conjugates, Adv. Drug Deliv. Rev., 56(8), 

1177–1192, 2004. 

12. Gabizon, A., Horowitz, A. T., Goren, D., Tzemach, D., Shmeeda, H., and Zalipsky, S., In vivo fate of 

folate-targeted polyethylene-glycol liposomes in tumor-bearing mice, Clin. Cancer Res., 9(17), 

6551–6559, 2003. 

13. Pan, X. and Lee, R. J., Tumour-selective drug delivery via folate receptor-targeted liposomes, Expert 

Opin. Drug Deliv., 1(1), 7–17, 2004. 

14. Campbell, R. B., Fukumura, D., Brown, E. B., Mazzola, L. M., Izumi, Y., Jain, R. K., Torchilin, V. P., 

and Munn, L. L., Cationic charge determines the distribution of liposomes between the vascular and 

extravascular compartments of tumors, Cancer Res., 62(23), 6831–6836, 2002. 

15. Sethi, V., Onyuksel, H., and Rubinstein, I., Liposomal vasoactive intestinal peptide, Methods 

Enzymol., 391, 377–395, 2005. 

16. Dubey, P. K., Mishra, V., Jain, S., Mahor, S., and Vyas, S. P., Liposomes modified with cyclic RGD 

peptide for tumor targeting, J. Drug Target., 12(5), 257–264, 2004. 

17. Gupta, B., Levchenko, T. S., and Torchilin, V. P., Intracellular delivery of large molecules and small 

particles by cell-penetrating proteins and peptides, Adv. Drug Deliv. Rev., 57(4), 637–651, 2005. 

18. Luo, Y. and Prestwich, G. D., Cancer-targeted polymeric drugs, Curr. Cancer Drug Targets, 2(3), 

209–226, 2002. 

19. Luo, Y., Bernshaw, N. J., Lu, Z. R., Kopecek, J., and Prestwich, G. D., Targeted delivery of doxorubicin 

by HPMA copolymer-hyaluronan bioconjugates, Pharm. Res., 19(4), 396–402, 2002. 

20. Mitra, A., Nan, A., Papadimitriou, J. C., Ghandehari, H., and Line, B. R., Polymer–peptide conjugates 

for angiogenesis targeted tumor radiotherapy, Nucl. Med. Biol., 33(1), 43–52, 2006. 

21. Wen, X., Jackson, E. F., Price, R. E., Kim, E. E., Wu, Q., Wallace, S., Charnsangavej, C., Gelovani, 

J. G., and Li, C., Synthesis and characterization of poly(L-glutamic acid) gadolinium chelate: a new 

biodegradable MRI contrast agent, Bioconjug. Chem., 15(6), 1408–1415, 2004. 

22. Wang, X., Feng, Y., Ke, T., Schabel, M., and Lu, Z. R., Pharmacokinetics and tissue retention of 

(Gd-DTPA)-cystamine copolymers, a biodegradable macromolecular magnetic resonance imaging 

contrast agent, Pharm. Res., 22(4), 596–602, 2005. 

23. Zong, Y., Wang, X., Goodrich, K. C., Mohs, A. M., Parker, D. L., and Lu, Z. R., Contrast-enhanced 

MRI with new biodegradable macromolecular Gd(III) complexes in tumor-bearing mice, Magn. 

Reson. Med., 53(4), 835–842, 2005. 

24. Brannon-Peppas, L. and Blanchette, J. O., Nanoparticle and targeted systems for cancertherapy, Adv. 

Drug Deliv. Rev., 56(11), 1649–1659, 2004. 

25. Labhasetwar, V., Nanotechnology for drug and gene therapy: the importance of understanding molecular 

mechanisms of delivery, Curr. Opin. Biotechnol., 16(6), 674–680, 2005. 

26. Sahoo, S. K. and Labhasetwar, V., Enhanced antiproliferative activity of transferrin-conjugated 

paclitaxel-loaded nanoparticles is mediated via sustained intracellular drug retention, Mol. Pharm., 

2(5), 373–383, 2005. 

27. Kommareddy, S., Tiwari, S. B., and Amiji, M. M., Long-circulating polymeric nanovectors for tumorselective 

gene delivery, Technol. Cancer Res. Treat., 4(6), 615–625, 2005. 


10 


28. Kommareddy, S. and Amiji, M., Preparation and evaluation of thiol-modified gelatin nanoparticles for 

intracellular DNA delivery in response to glutathione, Bioconjug. Chem., 16(6), 1423–1432, 2005. 

29. Kaul, G. and Amiji, M., Tumor-targeted gene delivery using poly(ethylene glycol)-modified gelatin 

nanoparticles: in vitro and in vivo studies, Pharm. Res., 22(6), 951–961, 2005. 

30. Farokhzad, O. C., Jon, S., Khademhosseini, A., Tran, T. N., Lavan, D. A., and Langer, R., Nanoparticle–aptamer 

bioconjugates: a new approach for targeting prostate cancer cells, Cancer Res., 64(21), 

7668–7672, 2004. 

31. Farokhzad, O. C., Khademhosseini, A., Jon, S., Hermmann, A., Cheng, J., Chin, C., Kiselyuk, A., 

Teply, B., Eng, G., and Langer, R., Microfluidic system for studying the interaction of nanoparticles 

and microparticles with cells, Anal. Chem., 77(17), 5453–5459, 2005. 

32. Torchilin, V. P., Lipid-core micelles for targeted drug delivery, Curr. Drug Deliv., 2(4), 319–327, 

2005. 

33. Zeng, F., Liu, J., and Allen, C., Synthesis and characterization of biodegradable poly(ethylene glycol)block-poly(5-benzyloxy-trimethylene 

carbonate) copolymers for drug delivery, Biomacromolecules, 

5(5), 1810–1817, 2004. 

34. Roby, A., Erdogan, S., and Torchilin, V. P., Solubilization of poorly soluble PDT agent, mesotetraphenylporphin, 

in plain or immunotargeted PEG-PE micelles results in dramatically improved cancer 

cell killing in vitro, Eur. J. Pharm. Biopharm., 62(3), 235–240, 2006. 

35. Aliabadi, H. M. and Lavasanifar, A., Polymeric micelles for drug delivery, Expert Opin. Drug Deliv., 

3(1), 139–162, 2006. 

36. Rapoport, N., Combined cancer therapy by micellar-encapsulated drug and ultrasound, Int. J. Pharm., 

277(1,2), 155–162, 2004. 

37. Gao, Z. G., Lee, D. H., Kim, D. I., and Bae, Y. H., Doxorubicin loaded pH-sensitive micelletargeting 

acidic extracellular pH of human ovarian A2780 tumor in mice, J. Drug Target., 13(7), 391–397, 

2005. 

38. Jeong, J. H., Kim, S. H., Kim, S. W., and Park, T. G., In vivo tumor targeting of ODN-PEG-folic 

acid/PEI polyelectrolyte complex micelles, J. Biomater. Sci. Polym. Ed., 16(11), 1409–1419, 2005. 

39. Ambade, A. V., Savariar, E. N., and Thayumanavan, S., Dendrimeric micelles for controlled drug 

release and targeted delivery, Mol. Pharm., 2(4), 264–272, 2005. 

40. Lee, J. H., Lim, Y. B., Choi, J. S., Lee, Y., Kim, T. I., Kim, H. J., Yoon, J. K., Kim, K., and Park, J. S., 

Polyplexes assembled with internally quaternized PAMAM-OH dendrimer and plasmid DNA have a 

neutral surface and gene delivery potency, Bioconjug. Chem., 14(6), 1214–1221, 2003. 

41. Bhadra, D., Bhadra, S., Jain, P., and Jain, N. K., Pegnology: a review of PEG-ylated systems, 

Pharmazie, 57(1), 5–29, 2002. 

42. Asthana, A., Chauhan, A. S., Diwan, P. V., and Jain, N. K., Poly(amidoamine) (PAMAM) dendritic 

nanostructures for controlled site-specific delivery of acidic anti-inflammatory active ingredient, 

AAPS Pharm. Sci. Tech., 6(3), E536–E542, 2005. 

43. Khan, M. K., Nigavekar, S.S, Mine, L. D., Kariapper, M. S., Nair, B. M., Lesniak, W. G., and Balogh, 

L. P., In vivo biodistribution of dendrimers and dendrimer nanocomposites: Implications for cancer 

imaging and therapy, Technol. Cancer Res. Treat., 4(6), 603–613, 2005. 

44. Lesniak, W., Bielinska, A. U., Sun, K., Janczak, K. W., Shi, X., Baker, J. R., and Balogh, L. P., 

Silver/dendrimer nanocomposites as biomarkers: Fabrication, characterization, in vitro toxicity, and 

intracellular detection, Nano Lett., 5(11), 2123–2130, 2005. 

45. D’Souza, G. G., Boddapati, S. V., and Weissig, V., Mitochondrial leader sequence: Plasmid DNA 

conjugates delivered into mammalian cells by DQAsomes co-localized with mitochondria, Mitochondrion, 

5(5), 352–358, 2005. 

46. D’Souza, G. G., Rammohan, R., Cheng, S. M., Torchilin, V. P., and Weissig, V., DQAsomemediated 

delivery of plasmid DNA toward mitochondria in living cells, J. Controlled Release, 92(1,2), 

189–197, 2003. 


2 

CONTENTS 

Passive Targeting of Solid 

Tumors: Pathophysiological 

Principles and Physicochemical 

Aspects of Delivery Systems 

S. M. Moghimi 

2.1 Introduction ...........................................................................................................................11 

2.2 Barriers to Extravasation.......................................................................................................12 

2.3 Selected Delivery Systems....................................................................................................12 

2.3.1 Liposomes..................................................................................................................12 

2.3.2 Polymeric Nanoparticles ...........................................................................................14 

2.3.3 Nanotechnology-Derived Nanoparticles ...................................................................15 

2.3.4 Macromolecular and Related Delivery .....................................................................16 

2.4 Conclusions ...........................................................................................................................16 

References ......................................................................................................................................17 

2.1 INTRODUCTION 

A solid tumor comprises two major cellular components: the tumor parenchyma and the stroma; the 

latter incorporating the vasculature and other supporting cells. As the tumor grows, in order to meet 

the metabolic requirements of an expanding population of tumor cells, the pre-existing blood 

vessels become subject to intense angiogenic pressure. Several factors produced by tumor cells 

and infiltrating immune-competent effector cells in the tumor parenchyma are believed to signal the 

development of new capillaries from the pre-existing vessels by capillary sprouting and/or dysregulated 

intussusceptive microvascular growth. 1 Further, in many solid tumors, endothelial cells 

destined to create new vessels are recruited not only from nearby vessels, but also to a significant 

extent from precursor cells within the bone marrow (so-called endothelial progenitor cells), 

a process referred to as “vasculogenesis.” 2 

Scanning electron microscopy of microvascular corrosion casts has allowed visualization of the 

geometry of blood vessel architecture in solid tumors. From these studies, it has become apparent 

that tumor blood vessels are highly irregular and show gross architectural changes that differ from 

those in normal organs and from newly formed blood vessels, such as those found in wound healing 

and in other angiogenic sites. 1,3 For instance, the thickness of a tumor blood vessel wall is poorly 


11

12 

correlated to its diameter. Therefore, despite the large size of some tumor vessels, the tumor blood 

flows is chaotic, with high flow rates in some segments and stagnation in others. 1 Also, the blood 

flow may temporarily change direction within individual tumor vessels. 

Further, the structure and organization of the endothelial cells, pericytes, and vascular basement 

membrane of tumor vessels are all abnormal. 1,3–6 One consistent abnormality of tumor blood vessels 

is their high permeability to macromolecules, arising from irregularly shaped and loosely interconnected 

endothelial cells (where the size of fenestrae often ranges from 200 to 2000 nm) and their 

less frequent and intimate association with pericytes and the vascular basement membrane. 1,3–5 

Marked variability has been noted in endothelial permeability among different tumors, different 

vessels within the same tumor, and during tumor growth, regression, and relapse. The extent of 

tumor blood vessel permeability is also controlled by the host microenvironment, and increases with 

the histological grade and malignant potential of tumors. 7,8 

Given the potency and toxicity of modern pharmacological agents, tissue selectivity is a major 

issue. In the delivery of chemotherapeutic agents to solid tumors this is particularly critical, 

since the therapeutic window for these agents is often small and the dose–response curve steep. 

Therefore, the idea of exploiting the well-documented vascular abnormalities of tumors, restricting 

penetration into normal tissue interstitium while allowing freer access to that of the tumor, becomes 

particularly attractive. 

2.2 BARRIERS TO EXTRAVASATION 

As a consequence of temporal and spatial heterogeneity in tumor blood flow, solid tumors usually 

contain well-perfused, rapidly growing regions, and poorly perfused, often necrotic areas. 1,3 As in 

normal tissues, diffusive and convective forces govern the movement of molecules into the interstitium 

of tumors. However, diffusion is believed to play a minor role in the movement of solutes 

across the endothelial barrier in comparison with bulk fluid flow. Examination of pressure gradients 

in experimental tumors has suggested that the movement of macromolecules and particulate 

materials out of the tumor blood vessels and into the extra-vascular compartment is remarkably 

limited. This has been attributed to a higher-than-expected interstitial pressure, in part due to a lack 

of functional lymphatic drainage, coupled with lower intra-vascular pressure. 3 In addition, interstitial 

pressure tends to be higher at the center of solid tumors, diminishing towards the periphery, 

creating a mass flow movement of fluid away from the central region of the tumor. 3 For example, 

the measured interstitial fluid pressure in invasive breast ductal carcinoma was 29G3 mm Hg, 

compared with 3.0G0.8 mm Hg in patients with benign tumors and K3.0G0.1 mm Hg in patients 

with normal breast parenchyma. 9 Nevertheless, the lower interstitial pressure in the periphery still 

permits adequate extravasation of fluid and macromolecules. 

These pathophysiological characteristics have serious implications for the systemic delivery of 

not only low-molecular-weight and macromolecular agents, but also particulate delivery vehicles. 

Simply enhancing the plasma half-life of these agents (e.g., long-circulating carriers) will not 

necessarily lead to an increase in therapeutic effect. 10 Furthermore, distribution, organization, 

and relative levels of collagen, decorin, and hyaluronan also impede the diffusion of extravasated 

macromolecules and particulate systems in tumors. 11 Thus, diffusion of macromolecules and 

particles will vary with tumor types, anatomical locations, and possibly by factors that influence 

extracellular matrix composition and/or structure. 

2.3 SELECTED DELIVERY SYSTEMS 

2.3.1 LIPOSOMES 


Liposomes are perhaps the best studied vehicles in cancer drug delivery, capable of either 

increasing the drug concentration in solid tumors and/or limiting drug exposure to critical target 


Passive Targeting of Solid Tumors 13 

sites such as bone marrow and myocardium. 10,12 For example, Myocete is a liposomal formulation 

of doxorubicin (an inhibitor of topoisomerase II) approximately 190 nm in size that was approved 

by the European Agency for the Evaluation of Medicinal Products (EMEA) in 2000 for the treatment 

of metastatic breast cancer. This formulation provides a limited degree of prolonged 

circulation when compared with doxorubicin in the free form. Myocete releases more than half 

of its associated doxorubicin within a few hours of administration and 90% within 24 h. Similar to 

intravenously injected nanoparticulate systems, liposomes are rapidly intercepted by macrophages 

of the reticuloendothelial system. 10 Hepatic deposition of Myocete could lead to gradual release of 

the cytotoxic agent back to the systemic circulation (a macrophage depot system), as well as 

induction of Kupffer cell apoptosis. 10 Following apoptosis, restoration of Kupffer cells may take 

up to two weeks. 13 A potentially harmful effect is the occurrence of bacteriemia during the period of 

Kupffer cell deficiency. Although Myocete administration decreases the frequency of cardiotoxicity 

and neutropenia compared with free drug, 14 there is still controversy as to whether liposomal 

encapsulation exhibits equivalent efficacy to doxorubicin. 15 

Macrophage deposition of intravenously administered liposomes can be markedly minimized 

either by bilayer or surface modification. 10,16 Regulatory approved examples include DaunoXome w , 

a daunorubicin-encapsulated liposome 45 nm in size with a rigid bilayer for HIV-related Kaposi’s 

sarcoma, and Doxil w /Caelyx w , a poly(ethylene glycol)-grafted rigid vesicle of 100-nm diameter 

with encapsulated doxorubicin for HIV-related Kaposi’s sarcoma and refractory ovarian carcinoma. 

As a result of their small size, rigid bilayer, and hydrophilic surface display (as in the case of 

Doxil w ), these formulations exhibit poor surface opsonization, a process that limits vesicle recognition 

by macrophages in contact with the blood and consequently prolongs their residency time 

within the vasculature. 10,16 For instance, Doxil w has a biphasic circulation half-life of 84 min and 

46 h in humans. In addition, Doxil w also has a high drug loading capacity; here doxorubicin is loaded 

actively by an ammonium sulfate gradient (as doxorubicin sulfate) yielding liposomes with a high 

content of doxorubicin aggregates, which remain highly stable within the vasculature with minimum 

drug loss. 17 Therefore, it is not surprising to see that such liposomal formulations exhibit favorable 

pharmacokinetics when compared with the free drug. For example, the area under the curve after 

a dose of 50 mg/m 2 doxorubicin encapsulated in long-circulating liposomes is approximately 

300-fold greater than that of free doxorubicin. 17 In addition, clearance and volume of distribution 

are reduced by at least 250- and 60-fold, respectively. 17 However, as a result of their prolonged 

circulation times, alternative toxic reactions have been reported with such vehicles. The most 

notable dose-limiting toxicity associated with continuous infusion of Doxil w is palmar–plantar 

erythrodysesthesis. 17 

Following extravasation into solid tumors, long-circulating liposomes often distribute heterogeneously 

in perivascular clusters that do not move significantly and poorly interact with cancer cells. 18 

Therefore, the efflux of drug must follow the process of liposome extravasation at a rate that maintains 

free drug levels in the therapeutic range. The rate of drug release from liposomes not only depends 

on the composition of the interstitial fluid surrounding tumors but also on the drug type and encapsulation 

procedures. The importance of the latter is highlighted by the observation that extravasated 

long-circulating cisplatin-containing liposomes (where cisplatin is loaded passively) lack anti-tumor 

activity, whereas cisplatin in free form is capable of inserting cytotoxicity. 19 This is in contrast to the 

effective anti-tumor property of the same liposomal lipid composition containing entrapped doxorubicin. 

It is believed that nonspecific chemical disruption or collapse of the liposomal pH gradient, that 

is used to load liposomes actively with doxorubicin, may trigger doxorubicin release. 17 

Long-circulating liposomes have the capability to deliver between 3 and 10 times more drug to 

solid tumors compared with the administered drug in its free form. If the entrapped drugs are 

released from extravasated liposomes, it is very likely that these vesicles inherently overcome 

a certain degree of multidrug resistance by the tumor cells. Thus, tumor regression is to be expected 

with tumors exhibiting a low resistance factor. With tumors exhibiting higher resistance 

levels, due to over-expression of energy-dependent efflux pumps such as P-glycoprotein and 


14 

multidrug-resistance-related protein, alternative approaches are necessary. One effective strategy is 

to use long-circulating temperature-sensitive liposomes in conjugation with hyperthermia, but this 

approach has limited applicability for visceral and widespread malignancies. 20 Others have elaborated 

on biochemical triggers such as cleaveable poly(ethylene glycol)-phospholipid conjugates to 

generate fusion competent vesicles 21 and enzyme-mediated liposome destabilization and pore 

formation. 22–24 Examples of the latter include long-circulating liposomes with attached proteasesensitive 

haemolysin 22 and pro-drug ether liposomes (e.g., vesicles containing phospholipids with 

a nonhydrolyzable ether bond in the 1-position), which are susceptible to degradation by secretory 

phospholipases. 23,24 For instance, the level of secretory phospholipases (such as the secretory 

phospholipase A2) is dramatically elevated within the interstitium of various tumors. 

Secretory phospholiapse A2 not only acts as a trigger resulting in the release of encapsulated 

cytotoxic drugs from pro-drug ether liposomes, but also generates highly cytotoxic lysolipids 

that destabilizes plasma membrane of tumor cells, thereby enhancing their permeability to cytotoxic 

drugs. 24 

There are several approaches that exploit active targeting of long-circulating liposomes to 

tumor cells, where receptor-mediated internalization is strongly believed to bypass tumor cell 

multidrug-efflux pumps. 10,16,17,25 These strategies utilize tumor-specific monoclonal antibodies 

or their internalizing epitopes, or ligands, such as folic acid, which are attached to the distal end 

of the poly(ethylene glycol) chains expressed on the surface of long-circulating liposomes. Nevertheless, 

with such approaches the delivery part is still passive and relies on liposome extravasation. 

2.3.2 POLYMERIC NANOPARTICLES 


Abraxanee is the only example of a regulatory approved (FDA, USA) nanoparticle formulation for 

intravenous drug delivery in cancer patients. It is paclitaxel bound to albumin nanoparticles, with a 

mean diameter of 130 nm, for use in individuals with metastatic breast cancer who have failed 

combination chemotherapy or relapse within 6 months of adjuvant chemotherapy. This formulation 

overcomes poor solubility of paclitaxel in the blood and allows patients to receive 50% 

more paclitaxel per dose over a 30-min period. 26,27 Unlike Cremophor w EL/ethanol or Tween w 

80-solubilized taxanes, acute hypersensitivity reactions, which are secondary to complement 

activation, have yet to be reported following Abraxanee infusion. Albumin nanoparticles seem 

to interact with gp60 receptors present on tumor blood vessels that transport the nanoparticles into 

tumor interstitial spaces by transcytosis, a process that may partly contribute to the effectiveness of 

Abraxanee. However, hepatic deposition (Kupffer cell capture) and processing of a significant 

fraction of albumin nanoparticles are most likely to occur. Indeed, after a 30 min infusion 

of 260 mg/m 2 doses of Abraxanee, faecal excretion accounted for approximately 20% of the 

administered dose (ABRAXIS Oncology, A division of American Pharmaceutical Partners, Inc., 

Schaumburg, IL 60173, USA; 2005), thus supporting a role for hepatic handling and biliary 

excretion of albumin nanoparticles (or its components). 

Nanoparticles assembled from synthetic polymers have also received much attention in cancer 

drug delivery. 28 One interesting example is doxorubicin-loaded poly(alkyl cyanoacrylate) (PACA) 

nanoparticles. In vitro studies have indicated that PACA nanoparticles can overcome drug resistance 

in tumor cells expressing multidrug-resistance-1-type efflux pumps. 29 The mechanism of 

action is related to adherence of PACA nanoparticles to tumor cell plasma membrane, which 

initiates particle degradation and provides a concentration gradient for doxorubicin, and diffusion 

of doxorubicin across the plasma membrane following formation of an ion pair between the 

positively charged doxorubicin and the negatively charged cyanoacrylic acid (a nanoparticle 

degradation product). 29 These observations clearly indicate that drug release and nanoparticle 

degradation must occur simultaneously, yielding an appropriate size complex with correct physicochemical 

properties for diffusion across the plasma membrane. Further developments with 

PACA nanoparticles include preparations that contain doxorubicin within the particle core and 



cyclosphorin, an inhibitor of the P-glycoprotein, at the surface. 30 Similar to liposomes, long-circulating 

versions of PACA nanoparticles have also been engineered for passive as well as active 

targeting to solid tumors. 31 

2.3.3 NANOTECHNOLOGY-DERIVED NANOPARTICLES 

Nanotechnology is a cross-disciplinary field, which involves the ability to design and exploit the 

unique properties that emerge from man-made materials ranging in size from 1 to greater than 

100 nm. 16 Indeed, the physical and chemical properties of materials—such as porosity, electrical 

conductivity, light emission, and magnetism—can significantly improve or radically change 

as their size is scaled down to small clusters of atoms. These advances are beginning to have 

a paradigm-shifting impact not least in experimental (e.g., thermal tumor killing) and diagnostic 

oncology. 16,32 Examples include superparamagnetic iron oxide nanocrystals, quantum dots (QDs), 

inorganic nanoparticles, and composite nanoshells. The surfaces of these entities are amenable to 

modification with synthetic polymers (to afford long-circulating properties) and/or to targeting 

ligands. However, a key problem with these technologies is toxicity and is discussed elsewhere. 16 

Iron oxide nanocrystals are formed from an inner core of hexagonally shaped iron oxide 

particles of approximately 5 nm, which express correlated electron behavior; at a high enough 

temperature, they are superparamagnetic. 33,34 In addition, dextran or synthetic polymers such as 

poly(ethyleneglycol) surround the crystal core. Indeed, it is the combination of the small size and 

surface characteristics that allow iron oxide nanocrystals, once injected into the blood stream, to 

bypass rapid detection by the body’s defence cells and accumulate in tumor sites by extravasation. 

Therefore, they are useful for patient selection, detection of tumor progression, and tracking of the 

effectiveness of anti-tumor treatment regimens by magnetic resonance imaging (MRI). These 

approaches can be extended for site-specific imaging of tumor vasculature with targeting 

ligands. In addition, iron oxide nanocrystals can slowly extravasate from the vasculature into the 

interstitial spaces, from which they are transported to lymph nodes by way of lymphatic vessels. 34 

Within lymph nodes they are captured by local macrophages, and their intracellular accumulation 

shortens the spin relaxation process of nearby protons detectable by MRI. On magnetic resonance 

images, those node regions accumulating iron oxide appear dark relative to surrounding tissues. 

Indeed, iron oxide nanocrystals can distinguish between normal and tumor-bearing nodes and 

reactive and metastatic nodes. 34 

QDs are made of semiconductors like silicon and gallium arsenide. 35,36 In these particles there 

are discrete electronic energy levels (valance band and conduction band), but the spacing of the 

electronic energy levels (band gap) can be precisely controlled through variation in size. When 

a photon, with higher energy than the energy of the band gap, hits a QD, an electron is promoted 

from valance band into the conduction band, leaving a hole behind. Electrons emit their excess 

energy as light when they recombine with holes. Since optical response is due to the excitation 

of single electron-hole pairs, the size and shape of QDs can be tailored to fluoresce specific colors. 

The ability of QDs to tune broad wavelength together with their photostability is of paramount 

importance in biological labeling. 35,36 Indeed, QDs stay lit much longer than conventional dyes 

used for imaging and tagging purposes and therefore have the potential to improve the resolution of 

tumor cells to the single cell level by optical imaging as well as determining heterogeneity among 

cancer cells in a solid tumor. 37–39 Unlike QDs, where optical response is due to the excitation of 

single electron hole pairs, in metallic nanoparticles (e.g., gold) incident light can couple to the 

plasmon excitation of the metal. This involves the light-induced motion of all valence electrons. 36 

Therefore, the type of plasmon that exists on a surface of a metallic nanoparticle is directly related 

to its shape and curvature; so it is possible to make a wide range of light scatterers that can be 

detected at different wavelengths. Composite nanoshells consist of a spherical dielectric core (e.g., 

silica) surrounded by a thin metal shell (e.g., gold). Again, by controlling the relative thickness of 

the core and shell layers of the composite nanoparticle, the plasmon resonance and the resultant 


16 

optical absorption properties can be tuned from near-UV to the mid-infrared. Of particular interest 

is the ability of near-infrared light (700–1000 nm) to penetrate through tissue at depths of a few cm 

with minimal heat generation and tissue damage. Thus, a recent study demonstrated rapid irreversible 

photothermal ablation of tumor tissue in vivo following administration of near-infraredabsorbing 

silica–gold nanoshells in combination with an extracorporeal low-power diode laser. 40 

2.3.4 MACROMOLECULAR AND RELATED DELIVERY 

Polymer-based drug delivery systems also favorably alter the pharmacokinetics and biodistribution 

of conjugated drugs and accumulate in tumor interstitium following extravasation. 41 Examples 

include SMANCS (a conjugate of the polymer styrene-co-maleic acid/anhydride and neocarzinostatin 

for treatment of hepatocellular carcinoma), conjugates of various cytotoxic agents (e.g., 

paclitaxel, doxorubicin, platinate, and campthothecin) with polyglutamate and nonbiodegradable 

hydroxypropyl methacrylamide. 

Other related polymer-based systems in cancer drug delivery include micelles 42,43 and dendrimers. 

44 For example, Pluronics w (copolymers of ethylene oxide and propylene oxide) are capable 

of forming micelles, and some members of Pluronic copolymers can overcome multidrug resistance. 

42 However, it is becoming clear that Pluronic copolymers can induce complement activation, 

even at concentrations below their critical micelle concentration, which may increase the risk 

of pseudoallergy in sensitive patients. 45 

Dendrimers are highly branched macromolecules with controlled near monodisperse threedimensional 

architecture emanating from a central core. 44 Polymer growth starts from a central 

core molecule and growth occurs in an outward direction by a series of polymerization reactions. 

Hence, precise control over size can be achieved by the extent polymerization, starting from a few 

nanometers. Cavities in the core structure and folding of the branches create cages and channels. 

The surface groups of dendrimers are amenable to modification and can be tailored for specific 

applications. Therapeutic and diagnostic agents are usually attached to surface groups on dendrimers 

by chemical modification. For example, a recent study has used tagged-dendrimers for in vivo 

evaluation of tumor-associated matrix metalloproteinase-7 (matrilysin) activity. 46 

Other macromolecular systems for cancer targeting and treatment include various forms of 

monoclonal and bispecific monoclonal antibodies against tumor-associated antigens. 47 These can 

further be coupled to drugs, toxins, enzymes (as in antibody-directed enzyme pro-drug therapy), 

cytokines, radionuclides, etc. 

2.4 CONCLUSIONS 


The chaotic blood flow in tumor vasculature and the heterogeneous vascular permeability of tumor 

blood vessels are among the key barriers controlling passive delivery of macromolecular and 

particulate delivery systems into the interstitium of solid tumors. Already compromised by 

abnormal hydrostatic pressure gradients, compressive mechanical forces generated by tumor cell 

proliferation cause intratumoral vessels to compress and collapse, thus creating further barriers for 

passive targeting. Interestingly, tumor-specific cytotoxic therapy, reducing tumor cell number, may 

result in more efficient delivery by decompressing these same vessels, 48 but this enhanced perfusion 

could provide a route for metastasis. Pathohysiologoical barriers, however, are not fully developed 

in micrometastases, and also pose a lesser problem in the diagnostic oncology as well as in drug 

delivery to well-perfused and low-pressure regions in larger tumors. Some of the problems may 

possibly be overcome by design of long-circulating multifunctional carriers (carriers that contain 

appropriate combinations of cytotoxic agents, diagnostic, and barrier-avoiding components) with 

biochemical triggering mechanisms. 16 The vascular barrier of the solid tumor is also its Achilles’ 

heel; the nutritionally demanding tumor cells are entirely dependent upon a functional vasculature. 



For this reason, interest has also been focused on the concept of tumor vasculature as a target rather 

than a barrier and is reviewed in this compendium and elsewhere. 49–51 

REFERENCES 

1. Munn, L. L., Aberrant vascular architecture in tumors and its importance in drug-based therapies, 

Drug Discov. Today, 8, 396, 2003. 

2. Lyden, D. et al., Impaired recruitment of bone-marrow-derived endothelial and haematopoietic 

precursor cells blocks, tumor angiogenesis, and growth, Nat. Med., 7, 1194, 2001. 

3. Jain, R. K., Delivery of molecular medicine to solid tumors: Lessons from in vivo imaging of gene 

expression and function, J. Control. Release, 74, 7, 2001. 

4. Dvorak, H. F. et al., Identification and characterization of the blood vessels of solid tumors that are 

leaky to circulating macromolecules, Am. J. Pathol., 133, 95, 1988. 

5. Hashizume, H. et al., Openings between defective endothelial cells explain tumor vessel leakiness, 

Am. J. Pathol., 156, 1363, 2000. 

6. Morikawa, S. et al., Abnormalities in pericytes on blood vessels and endothelial sprouts in tumors, 

Am. J. Pathol., 160, 985, 2002. 

7. Fukumura, D. et al., Effect of host microenvironment on the microcirculation of human colon adenocarcinoma, 

Am. J. Pathol., 151, 679, 1997. 

8. Daldrup, H. et al., Correlation of dynamic contrast-enhanced MR imaging with histologic tumor 

grade: Comparison of macromolecular and small-molecular contrast media, Am. J. Roentgenol., 

171, 941, 1998. 

9. Nathanson, S. D. and Nelson, L., Interstitial fluid pressure in breast cancer, benign breast conditions, 

and breast parenchyma, Ann. Surg. Oncol., 1, 333, 1994. 

10. Moghimi, S. M., Hunter, A. C., and Murray, J. C., Long-circulating and target-specific nanoparticles: 

Theory to practice, Pharmacol. Rev., 53, 283, 2001. 

11. Pluen, A. et al., Role of tumor–host interactions in interstitial diffusion of macromolecules: Cranial vs. 

subcutaneous tumors, Proc. Natl Acad. Sci. USA, 98, 4628, 2001. 

12. Torchilin, V. P., Recent advances with liposomes as pharmaceutical carriers, Nat. Rev. Drug Discov., 

4, 145, 2005. 

13. Daemen, T. et al., Liposomal doxorubicin-induced toxicity: Depletion and impairment of phagocytic 

activity of liver macrophages, Int. J. Cancer, 61, 716, 1995. 

14. Batist, G. et al., Reduced cardiotoxicity and preserved antitumor efficacy of liposome-encapsulated 

doxorubicin and cyclophosphamide compared with conventional doxorubicin and cyclophosphamide 

in a randomized, multicenter trial of metastatic breast cancer, J. Clin. Oncol., 19, 1444, 2001. 

15. Williams, G., Cortazar, P., and Pazdur, R., Developing drugs to decrease the toxicity of 

chemotherapy, J. Clin. Oncol., 19, 3439, 2001. 

16. Moghimi, S. M., Hunter, A. C., and Murray, J. C., Nanomedicine: Current status and future prospects, 

FASEB J., 19, 311, 2005. 

17. Gabizon, A., Shmeeda, H., and Barenholz, Y., Pharmacokinetics of pegylated liposomal doxorubicin: 

Review of animal and human studies, Clin. Pharmacokinet., 42, 419, 2003. 

18. Yuan, F. et al., Microvascular permeability and interstitial penetration of sterically stabilized (stealth) 

liposomes in a human tumor xenograft, Cancer Res., 54, 3352, 1994. 

19. Bandak, S. et al., Pharmacological studies of cisplatin encapsulated in long-circulating liposomes in 

mouse tumor models, Anti-Cancer Drugs, 10, 911, 1999. 

20. Huang, S. K. et al., Liposomes and hyperthermia in mice: Increased tumor uptake and therapeutic 

efficacy of doxorubicin in sterically stabilized liposomes, Cancer Res., 54, 2186, 1994. 

21. Zalipsky, S. et al., New detachable poly(ethylene glycol) conjugates: Cysteine–cleavable lipopolymers 

regenerating natural phospholipid, diacylphosphatidylethanolamine, Bioconjug. Chem., 10, 703, 

1999. 

22. Provoda, C. J. and Lee, K.-D., Bacterial pore-forming hemolysins and their use in the cytosolic 

delivery of macromolecules, Adv. Drug Deliv. Rev., 41, 209, 2000. 

23. Andresen, T. L., Jensen, S. S., and Jørgensen, K., Advanced strategies in liposomal cancer therapy: 

Problems and prospects of active and tumor specific drug release, Prog. Lipid Res., 44, 68, 2005. 


18 


24. Andresen, T. L. et al., Triggered activation and release of liposomal prodrugs and drugs in cancer 

tissue by secretory phospholipase A2, Curr. Drug Deliv., 2, 353, 2005. 

25. Allen, T. M., Ligand-targeted therapeutics in anti-cancer therapy, Nat. Rev. Cancer, 2, 750, 2002. 

26. Bartels, C. L. and Wilson, A. F., How does a novel formulation of paclitaxel affect drug delivery in 

metastatic breast cancer?, US Pharm., 29, HS18, 2004. 

27. Garber, K., Improved paclitaxel formulation hints at new chemotherapy approach, J. Natl Cancer 

Inst., 96, 90, 2004. 

28. Brigger, I., Dubernet, C., and Couvreur, P., Nanoparticles in cancer therapy and diagnosis, Adv. Drug 

Deliv. Rev., 54, 631, 2002. 

29. Colin de Verdiere, A. et al., Reversion of multidrug resistance with polyalkylcyanoacrylate nanoparticles: 

Towards a mechanism of action, Br. J. Cancer, 76, 198, 1997. 

30. Soma, C. E. et al., Reversion of multidrug resistance by co-encapsulation of doxorubicin and cyclosporin 

A in polyalkylcyanoacrylate nanoparticles, Biomaterials, 21, 1, 2000. 

31. Stella, B. et al., Design of folic acid-conjugated nanoparticles for drug targeting, J. Pharm. Sci., 89, 

1452, 2000. 

32. Ferrari, M., Cancer nanotechnology: Opportunities and challenges, Nat. Rev. Cancer, 5, 161, 2005. 

33. Perez, J. M., Josephson, L., and Weissleder, R., Use of magnetic nanoparticles as nanosensors to probe 

for molecular interactions, Chem. Bio. Chem., 5, 261, 2004. 

34. Moghimi, S. M. and Bonnemain, B., Subcutaneous and intravenous delivery of diagnostic agents to 

the lymphatic system: Applications in lymphoscintigraphy and indirect lymphography, Adv. Drug 

Deliv. Rev., 37, 295, 1999. 

35. Medintz, I. L. et al., Quantum dot bioconjugates for imaging, labelling, and sensing, Nat. Mater., 

4, 435, 2005. 

36. Alivisatos, P., The use of nanocrystals in biological detection, Nat. Biotechnol., 22, 47, 2003. 

37. Stroh, M. et al., Quantum dots spectrally distinguish multiple species within the tumor milieu in vivo, 

Nat. Med., 11, 678, 2005. 

38. Wu, X. et al., Immunofluorescent labeling of cancer marker Her2 and other cellular targets with 

semiconductor quantum dots, Nat. Biotechnol., 21, 41, 2003. 

39. Gao, X. et al., In vivo cancer targeting and imaging with semiconductor quantum dots, Nat. 

Biotechnol., 22, 969, 2004. 

40. Hirsch, L. R. et al., Nanoshell-mediated near-infrared thermal therapy of tumors under magnetic 

resonance guidance, Proc. Natl Acad. Sci. USA, 100, 13549, 2003. 

41. Duncan, R., The dawning era of polymer therapeutics, Nat. Rev. Drug Discov., 2, 347, 2003. 

42. Oh, K. T., Bronich, T. K., and Kabanov, A. V., Micellar formulations for drug delivery based on 

mixtures of hydrophobic and hydrophilic Pluronic w block copolymers, J. Control. Release, 94, 411, 

2004. 

43. Adams, M. L., Lavasanifar, A., and Kwon, G. S., Amphiphilic block copolymers for drug delivery, 

J. Pharm. Sci., 92, 1343, 2003. 

44. Haag, R., Supramolecular drug-delivery systems based on polymeric core-shell architectures, Angew. 

Chem. Int. Ed., 43, 278, 2004. 

45. Moghimi, S. M. et al., Causative factors behind poloxamer 188 (Pluronic F68 Flocor)-induced complement 

activation in human sera: a protective role against poloxamer-mediated complement activation 

by elevated serum lipoprotein levels, Biochim. Biophys. Acta-Mol. Basis Dis., 103, 1689, 2004. 

46. McIntyre, J. O. et al., Development of a novel fluorogenic proteolytic beacon for in vivo detection 

and imaging of tumor-associated matrix metalloproteinase-7 activity, Biochem. J., 377, 617, 2004. 

47. Kosterink, J. G. W. et al., Strategies for specific drug targeting to tumor cells. Vol. 12 of Drug Targeting, 

Organ-Specific Strategies, Molema, G., and Meijer, D. K. F., Eds., Wiley-VCH, Weinheim pp.199–232, 

2001, chapter 8. 

48. Padera, T. P. et al., Cancer cells compress intratumor vessels, Nature, 427, 695, 2004. 

49. Pasqualini, R., Arap, W., and McDonald, D. M., Probing the structural and molecular diversity of 

tumor vasculature, Trend Mol. Med., 8, 563, 2002. 

50. Murray, J. C. and Moghimi, S. M., Endothelial cells as therapeutic targets in cancer: New biology 

and novel delivery systems, Crit. Rev. Ther. Drug Carrier Syst., 20, 139, 2003. 

51. Reynolds, A. R., Moghimi, S. M., and Hodivala-Dilke, K., Nanoparticle-mediated gene delivery 

to tumor neovasculature, Trend Mol. Med., 9, 2, 2003. 


3 

CONTENTS 

Active Targeting Strategies in 

Cancer with a Focus on Potential 

Nanotechnology Applications 


3.1 Introduction ........................................................................................................................... 19 

3.2 Nanoparticle Characteristics ................................................................................................. 20 

3.2.1 Composition and Biocompatibility ........................................................................... 22 

3.2.2 Derivitization .............................................................................................................23 

3.2.3 Detection.................................................................................................................... 23 

3.3 Nanoparticle Targeting.......................................................................................................... 24 

3.3.1 Inherent Targeting ..................................................................................................... 24 

3.3.2 Complicating Aspects ............................................................................................... 25 

3.3.3 Safety Issues .............................................................................................................. 26 

3.4 Targeting to Cancer Cells ..................................................................................................... 27 

3.4.1 Cell Surface Properties.............................................................................................. 28 

3.4.2 Metabolic Properties ................................................................................................. 29 

3.5 Targeting to Tumors.............................................................................................................. 30 

3.5.1 Aberrant Vasculature................................................................................................. 30 

3.5.2 Metabolic Environment............................................................................................. 32 

3.6 Fate of Nanoparticles ............................................................................................................ 32 

3.6.1 Site of Initial Application ......................................................................................... 33 

3.6.2 Distribution to Specific Organs................................................................................. 34 

3.6.3 Intracellular Uptake and Fate.................................................................................... 34 

3.7 Conclusions ........................................................................................................................... 35 

References ...................................................................................................................................... 37 


Nanotechnology has recently become a buzzword in several scientific fields, including the area 

of drug delivery, 1 and a variety of nanomedicine opportunities have recently been reviewed. 2 The 

concept of nanoparticles as drug delivery vehicles is not new as reviews on the subject were published 

nearly 20 years ago. 3 That nanoparticles could be selectively targeted by coating with monoclonal 

antibodies provided an early, ground-breaking proof-of-concept that these materials might one day 

be effective tools for the diagnosis and treatment of cancers. At their inception, all new technologies 


19

20 

Targeting 

Imaging 

appear to have a plethora of possibilities because limitations have not yet become apparent. Subsequent 

studies that examine the limits of various applications then provide a menu of feasible 

applications. This menu can change as new developments of the technology are realized that 

allow one to address additional applications. This has certainly been the case for nanotechnology 

applications that involve active targeting to cancers for diagnostic and therapeutic purposes. 4 

Although nanoparticles have shown tremendous promise in facilitating the targeted delivery of 

therapeutics and diagnostics to cancers, the composition and size of the particles have inherent 

physical and chemical properties that can compromise their capability to localize to and/or treat 

cancers because of non-selective cell and tissue uptake. Therefore, active targeting to cancers using 

nanoparticles requires consideration of not only unique properties of the cancer that allow for 

specific targeting but also attention to issues that minimize non-selective delivery to uninvolved 

regions of the body. Methods to overcome functional barriers that limit uptake of materials or 

delivery to cancer cells must be also considered. Even with proper consideration of these issues, 

successful disposition and targeting to cancers is quite a challenge. Certainly, lack of consideration 

of these issues can dramatically increase the risk of serious negative outcomes, particularly when 

targeted nanoparticles contain potent cytotoxic agents. 

Advantages and disadvantages of specific nanoparticles as well as methods to potentially 

correct shortcomings of nanoparticle targeting to cancers will be discussed in this chapter. 

Several of the topics raised in this chapter will also be discussed in much greater depth in other 

chapters in this text. Specifically, two major strategies of cancer targeting related to nanotechnology 

opportunities will be addressed: targeting to cancer cells and targeting to tumors. Attention will be 

paid to similarities as well as differences for these two strategies. Targeting cancer cells and tumors 

for diagnostic purposes as well as therapy will also be examined. Several applications for nanoparticles 

have been outlined by the National Cancer Institute (http://nano.cancer.gov) regarding 

unmet medical needs in the areas of targeting, imaging, delivery, and reporting agents for cancer 

diagnosis and treatment (Figure 3.1). General cellular responses to and fates of nanoparticles that 

can affect these potential applications will be discussed as critical aspects of ultimate 

clinical success. 

3.2 NANOPARTICLE CHARACTERISTICS 

Delivery 

Reporting 


FIGURE 3.1 Integrated potential applications of nanoparticles as visualized by the National Cancer Institute 

of the NIH. Nanoparticles, because of their general capacity for multiple modifications as well as their inherent 

properties, can have multiple functions relevant to targeting cancers for therapeutic and/or diagnostic purposes. 

Clinical success of nanoparticle-based diagnostics and therapeutics requires proper matching of 

particle characteristics. The characteristics of nanoparticles are critically dependent upon the 

materials used to prepare the nanoparticle. Nanoparticles can now be readily prepared from a 

wide range of inorganic and organic materials in a range of sizes from two to several hundred 

nanometers (nm) in diameter. Put in perspective, human cells are typically 10,000–20,000 nm in 

diameter. The plasma membrane of these cells is 6 nm in thickness. In most cases, nanoparticles 

can be generated to have narrow and defined size ranges. Other chapters in this text will focus on the 

physical and chemical characteristics of nanoparticles made from various materials as well as 


Active Targeting Strategies in Cancer with a Focus on Potential Nanotechnology Applications 21 

methods for their production. Although nanoparticles can be prepared from a wide variety of 

materials (inorganic salts, lipids, synthetic organic polymers, polymeric forms of amino acids, 

nucleic acids, etc.), this chapter will primarily focus on those prepared from materials that would 

be considered sufficiently safe for repeated systemic administrations and/or would be perceived to 

have an acceptable safety profile that would warrant use in man. In general, it is desirable for 

nanoparticles to be either readily metabolized or sufficiently broken down to produce only nontoxic 

metabolites that can be safely excreted. Indeed, tremendous advances have been made in 

controlling the chemical nature, degradable characteristics, and dimensions of nanoparticles. 

Many of the initial studies examining nanoparticles as delivery tools used particles prepared 

from materials such as polyalkylcyanoacrylates (PAA). 5 The extreme stability of PAA is both a 

positive and a negative. PAA nanoparticles will not be degraded prior to reaching a tissue or cell 

target site; however, once they reach that site, it is unlikely that they will be efficiently metabolized. 

Therefore, PAA nanoparticles have been extremely useful for initial studies of nanomaterials for 

cancer targeting, but an inability to clear PAA nanoparticles presents uncertainties as to their 

ultimate toxicological fate. Concerns over repeated PAA nanoparticle administrations in man 

and the need for more acceptable materials were highlighted early on. 3 One of the biggest concerns 

regarding poorly metabolized nanoparticles is that of accumulation and the potential sequelae 

associated with such an outcome. In some cases where a limited number of exposures would 

occur, one could consider the use of materials that are not readily metabolized by the body. In 

the case of certain cancer applications, it might be possible to use materials that otherwise would be 

considered to have an unacceptable safety signal following repeat dosing or that have the potential 

to accumulate. Therefore, rationales exist for the potential application of nanoparticles prepared 

from a wide range of materials, even those that, at first glance, would be considered unacceptable. 

Methods of production and composition define nanoparticle characteristics; these characteristics 

define potential issues (and opportunities) related to biocompatibility, derivitization, and 

detection. Nanoparticles can be prepared from a singular subunit that is chemically coupled and 

organized in a defined (e.g., dendrimers) or in a more random (e.g., polylactic acid) manner. 

Although these materials would not have a defined core, they can be impregnated with compatible 

materials and/or chemically modified at their surface. Materials such as glyconanoparticles would 

provide one approach where a distinct core with radiating ligands could be positioned using linkers. 

In such a case, the solid core, used to anchor each linker used for the attachment of targeting 

ligands, could be used to deliver a therapeutic or diagnostic payload. Liposomes are an example of 

nanoshell structures that can be loaded internally as well as impregnated within the shell. Many 

types of nanomaterials fall into one of these three general structural architectures (Figure 3.2). 

(a) (b) (c) 

FIGURE 3.2 General schema for three types of nanoparticle structures. (a) Nanoparticles can be formed from 

one type of material that can be impregnated with therapeutic or lipophilic imaging reagents (open diamonds) 

and modified with targeting ligands (crescents) positioned by chemical coupling through linker moieties. (b) 

Metal (or similar) cores (circles can be modified through a linker-targeting ligand system to generate another 

type of nanoparticle structure. In this case, it might be possible to use elaborated linkers as an environment 

compatible for incorporation of therapeutic or imaging reagents. (c) Shell-type nanoparticles such as liposomes 

where an aqueous compartment is enclosed by a bilayer of phospholipids can also be used for the targeted 

delivery of hydrophilic therapeutic or imaging reagents (filled hexagons). 


22 

3.2.1 COMPOSITION AND BIOCOMPATIBILITY 


Biological organic polymers can be formed from amino acids to form peptides and proteins. Some 

proteins have, by themselves, been used as nanoparticle delivery systems. Indeed, the protein 

ferritin functions as a coated nanoparticle. This molecule is w12 nm in diameter and can carry 

hydrous ferrous oxide (w5–7 nm in diameter) during its role as an iron storage system for the 

body. 6 Alternately, proteins can be used to generate nanoparticles for carrier applications; gelatin 

has been used to prepare nanoparticles. 7 Albumin nanoparticles have been described. 8 A 13-MDa 

ribonucleoprotein, termed a vault, has also been identified as a nanomaterial that could be used to 

deliver therapeutic and diagnostic agents. 9 Materials prepared from nucleic acids also have the 

potential to act as nanoparticle carriers. In general, nanoparticles prepared from biological materials 

would be biocompatible as a result of obvious elimination mechanisms. 

Nanoparticles can also be prepared from biological materials that are found in the body but 

are not typically organized as nanoparticle-size polymers; complex mixtures of polysaccharides, 

poly-lysine, and poly(D,L-lactic and glycolic acids; PLGA) have been prepared in a variety of sizes 

and in formats that allow ligand coupling with targeting moieties as well as diagnostic or therapeutic 

agents. 10 Synthetic organic polymers such as PLGA have been used to produce resorbable 

sutures, providing nanoparticles that will produce a sustained release of its contents. PLGA nanoparticles 

have been used to deliver wild-type p53 protein to cancer cells. 11 Such polymers have 

been studied for the development of nanoparticle delivery vehicles. 12–14 Although such nanoparticles 

would be considered relatively safe because of their biocompatibility, particles prepared from 

these types of materials can initiate inflammatory responses at sites of accumulation or deposit. 15 

A number of new synthetic organic polymer materials are being examined for their capacity to 

generate nanoparticles useful for drug encapsulation and targeting. 16 In general, organic polymerbased 

nanoparticles have provided promising results in the field of cancer targeting for diagnostics 

and therapy with the added capability of sustained release in some cases. 17 

Dendrimer structures are prepared from a series of repetitive chemical steps that perpetually 

increase their size as additional shells are added. Varying the seed molecule can produce nanoparticle 

structures of spherical or more elongated shapes. Dendrimer-based nanoparticle structures 

are inherently different from other organic-based nanoparticles prepared as linear sequences 

of subunits (e.g., amino acids used to form proteins) or from subunits that can randomly branch 

(e.g., polysaccharides). With the flexibility of chemistries that have now been described to prepare 

dendrimers, these materials show tremendous promise to prepare highly defined nanoparticles that 

can be targeted to cancers for therapeutic and diagnostic applications. To date, polyamidoamine 

(PAMAM) dendrimers have been the most extensively studied family of dendrimers that can be 

used to deliver anti-cancer drugs. 18 

A variety of inorganic materials can be used to generate nanoparticles for drug delivery that 

might be applied to cancer and tumor targeting. For example, iron oxide (Fe2O3) nanoparticles can 

be used to deliver anti-cancer agents. 19 Fe2O3 nanoparticles targeted to a cancer can become hot 

enough in an applied oscillating magnetic field to kill cells. Calcium phosphate precipitates can also 

be also made into nanoparticles. Although calcium phosphate precipitates can be metabolized over 

time and would be considered biocompatible, these materials can act as a potent adjuvant, potentially 

enhancing their application to target the delivery of cancer antigens. 20 In this way, calcium 

phosphate nanoparticles are similar to another inorganic salt precipitate, aluminum hydroxide 

(alum), that is currently approved as an adjuvant for human vaccines. 21 Semiconductor nanocrystals 

(quantum dots) are another example of inorganic nanoparticles. Quantum dots have exceptional 

characteristics for in vivo imaging and diagnostic applications. 22,23 However, some materials used 

to generate quantum dots such as CdSe can release toxic Cd 2C ions that alter ion channel function 

and lead to cell death when sufficient levels are reached. 24 Therefore, some of the inorganic 

materials used to generate nanoparticles may have significant biocompatibility issues. 



Lipids such as phospholipids and cholesterol can be used to generate single- or multi-lamellar 

spheres (liposomes) in the nanometer-size range. Liposome-based nanocapsules that can be loaded 

with diagnostic or therapeutic agents for the targeted delivery to cancers 25 and enhancement 

strategies for lipid-based nanoparticle-mediated tumor targeting have been reviewed. 26 Liposomes 

were some of the first nanoparticle structures extensively evaluated for cancer targeting. Early 

studies using liposomes highlighted issues associated with recognition and clearance by cells of the 

reticuloendothelial system (RES) that remove particulates from the systemic circulation. 27 Methods 

of masking liposomes from the RES such as modification with poly(ethylene glycol) (PEG) have 

been successfully used to limit RES clearance and increase circulating half-lives in serum. 28 

Solid lipid nanoparticles, nanostructure lipid carriers, and lipid–drug conjugate nanoparticles 

have also been described for the drug delivery strategies that could be applied to cancer diagnosis 

and/or therapy. 29 Lipid-based nanospheres can be sterically stabilized by the incorporation of 

artificial lipid derivatives that can be cross-linked. Subsequently, stabilized lipid-based nanospheres 

can be targeted to cancers using a conjugated antibody. 30 Such covalent modifications can improve 

the stability of lipid-based nanoparticles but can also reduce the biocompatible natures of these 

materials by modifying their clearance from the body. Apolipoprotein E-containing liposomes have 

also been prepared as a carrier for a lipophilic prodrug of daunorubicin as a means of targeting 

cancer cells that overexpress the receptor for low-density lipoproteins (LDL). 31 

3.2.2 DERIVITIZATION 

Because of their chemical and physical characteristics, nanoparticles exhibit inherent cellular 

targeting and uptake characteristics. Size and surface charge seem to be the two prominent characteristics 

that affect inherent nanoparticle targeting and cellular uptake. Because inherent targeting 

mechanisms may not provide the targeting or delivery characteristics desired, methods to modify 

nanoparticles with targeting agents can be critical. Although some nanoparticle materials are 

composed of materials with functional groups useful for chemical coupling, others are not. Such 

nanoparticle systems must be either modified to allow chemical coupling or doped with reagents 

that can be used for this modification. A number of coupling strategies have also been worked out 

that allow for efficient functionalization of nanomaterials through both reversible and irreversible 

chemistries. 14,32 These modifications allow for the coupling of antibodies, receptor ligands, and 

other potential targeting agents. Similar to the concerns associated with composition of the nanoparticle 

itself, any modification through chemical derivitization must also be considered with regard 

to generating materials with unacceptable toxicity or neutralization of the function of the nanoparticle 

or its targeting element. 

Nanoparticles have the advantage that they can be modified with multiple ligands to enhance 

their targeting selectivity and/or allow for simultaneous delivery of diagnostic and therapeutic 

agents. 33 It is important to appreciate the relative size of components used to construct and derivitize 

nanoparticles. For example, a quantum dot may be only 10 nm in diameter. Targeting that 

sized particle with an antibody might require the attachment of an IgG antibody that is roughly 

equal in size. By comparison, a fluorescent material that might be useful for localization of a 

targeted nanoparticle such as green fluorescent protein (GFP) is about 5 nm. Derivitization 

strategies for the construction of targeted nanoparticles must incorporate a consideration of 

potential steric conflicts for incorporation of targeting, detection, and therapeutic components. 

Other chapters in this text will extensively examine derivitization technologies for nanoparticles. 

3.2.3 DETECTION 

Most, if not all, nanoparticle structures currently investigated for delivery of cancer therapeutics 

also have the capacity to be detected or modified to contain a detectable agent that could be 

simultaneously used for cancer diagnosis. For example, PAMAM folate-dendrimers that contain 


24 

contrast media for detection by magnetic resonance imaging (MRI) are effectively targeted to 

cancer cells that overexpress the high affinity folate receptor. 34 Gadolinium ion–dendrimer nanoparticlesarereadilydetectedbyCTimagingand 

appear to provide several advantages over 

previous methodologies. 35 PAMAM dendrimer nanoparticles covalently coupled with a fluorescent 

label can be visualized as sites of increased retention. 36 Any strategy to produce nanoparticles that 

combines diagnostic and therapeutic elements, however, must consider potential conflicting 

aspects. Introduction of some heterocyclic anti-cancer molecules may act to quench fluorescent 

properties and the capacity to detect fluorescent labels. 

Some nanoparticles are particularly promising for cancer diagnosis because of their exceptional 

properties of detection using current radiographic and magnetic methods. Some new materials 

being prepared as nanoparticles will provide the potential for visualization using novel imaging 

methods that may lead to greater selectivity of signal and reduce false positives. 37 For example, 

quantum dots associated with metastatic cancer cells can be visualized using fluorescence emissionscanning 

microscopy. 38 Similarly, lipid-encapsulated liquid perfluorocarbon contrast media at the 

site of a tumor can be detected by ultrasonic acoustic transmission. 39 It is also possible to functionalize 

materials such as quantum dots to incorporate materials that can be photo-activated to 

enhance their activity or detection. 23 

3.3 NANOPARTICLE TARGETING 

Nanoparticles can be designed in a variety of ways to achieve targeted delivery. Some targeting 

strategies rely upon inherent properties of the particle, in particular, its composition, size, and 

surface properties. Furthermore, the particle itself can either be the agent being delivered, or it 

can be prepared to carry a cargo for delivery. Cargo release from the nanoparticles can occur while 

the nanoparticle is still relatively intact or through its decomposition. A number of methods have 

been described to integrate and retain cargo components within nanoparticles and these, in general, 

match to chemical or physical characteristics of the cargo with those of the material used to 

generate the nanostructure. For example, positively charged cargo can be held within the nanoparticle 

through interactions with an internal network such as a polyanionic polymer that resembles 

the organization of secretory granules synthesized by cells. 40 Alternately, organized complexes 

akin to coacervates proposed to participate in cell structure evolution can be formed between cargo 

and particle matrix. 40 Therefore, for some cancer-targeting strategies, one should consider not only 

compatibility of the nanoparticle carrier with its cargo but also degradation events that might affect 

temporal aspects of particle stability and cargo release. 

Some nanoparticles can be designed or delivered in such a way as to produce a default targeting 

event; other nanoparticles must be decorated on their surface to produce a targeted structure. 

Topical application of a nanoparticle system at the target site may be all that is required for a 

successful outcome. Such a simple approach is not typically sufficient for effective targeting of 

many cancers. Successful targeting may require reduction of inherent targeting tendencies for the 

material(s) used to prepare the nanoparticle. Depending upon the physical and chemical nature of 

the nanoparticle and the mode of administration, there can also be complicating factors that affect 

the effectiveness of the targeting method. Inherent targeting and complicating factors associated 

with some nanoparticles used for a targeted delivery can impart safety issues that must also be 

considered. With such characteristics, it is easy to see why active targeting of nanoparticles to 

cancers can be both complicated by competing biological events as well as facilitated by these same 

properties. 41 

3.3.1 INHERENT TARGETING 


The RES is composed of a series of sentinel cells located in several highly perfused organs, 

including the liver and spleen. 27 Nanoparticles can be rapidly cleared from the blood if they are 



recognized by RES cells in a non-selective fashion, typically before achievement of effective 

targeting. 13,42 In some instances, this inherent targeting can provide a means to selectively delivery 

materials. 43 In most cases, it is possible to modify the physical and chemical characteristics of 

nanoparticles to reduce their default uptake by the RES. 44 Methods to avoid the RES will be 

addressed in depth in other chapters in this text. In general, these measures follow principles 

initially outlined in the development of stealth liposomes that provided a means of extending the 

circulating half-life of a nanoparticle. Although PEG molecules of various lengths coupled using 

various chemistries 45 are frequently used in this approach, heparan sulfate glycosaminoglycans 

(HSGs) have also been shown to provide a protective coating that reduces immune detection. 46 

Interestingly, HSGs might be shed at tumors by tumor-associated heparanase activity. 

Another inherent targeting aspect of nanoparticles relates to the nature of tumor-associated 

vasculature. In general, nanoparticles smaller than 20 nm have the ability to transit out of blood 

vessels. Solid tumors grow rapidly; tumor-associated endothelial cells are continually bathed by a 

plethora of cancer cell-secreted growth factors. In turn, endothelial cells sprout new vessels to 

provide needed nutrients for the continued growth of the tumor. This cancer cell-endothelial cell 

relationship, however, leads to the establishment of a poorly organized vasculature that, under the 

constant drive of growth factor stimulation, fails to organize into a mature vascular bed. Therefore, 

tumor-associated vascular beds are poorly organized and more leaky that normal vasculature. 

Nanoparticles will inherently target to tumors as exudates through leaky vasculature. This phenomenon, 

referred to as the enhanced permeability and retention (EPR) effect, 47 will be covered 

extensively in other chapters in this text. 

Finally, peculiar surface properties of certain nanomaterials might affect their inherent 

interactions that could act to detract from a targeted delivery strategy. For example, some polyanionic 

dendrimers can be taken up by cells and act within those cells to interfere with replication of 

human immunodeficiency virus (HIV) that is considered to be the causative agent of AIDS. 48 

Although, from such studies, it is unclear if these dendrimers interact with the host cell or the 

pathogen to block their interaction; such a finding points to the potential for nanoparticles to 

interact with structures that might affect their cellular properties or cell function. In some cases, 

a nanoparticle with inherent capacity to interact with a cell or tissue might provide an added 

advantage of using that material for a specific indication. In other cases, such an inherent capacity 

to bind to or recognize cell or tissue components might highlight potential distractive aspects of that 

material for certain indications. 

3.3.2 COMPLICATING ASPECTS 

By their eponymous descriptor, nanoparticles have physical dimensions in the nanometer size scale 

similar to viruses and other materials that are either recognized by the body as pathogens or are 

elements associated with an infective event. Toll-like receptors (TLR) present on monocytes, 

leukocytes, and dendritic cells play a critical role in innate immunity with the capacity to recognize 

organized patterns present on viruses and bacteria. 49 TLR proteins are present on the surface of 

cells in the lung, spleen, prostate, liver, and kidney. Because the patterned surfaces of nanoparticles 

can look like pathogen components recognized by TLR proteins such as DNA, RNA, and repeating 

proteins like flagellin, it is possible that a number of cell types might non-selectively interact with 

some nanoparticles. If such an interaction occurs, there are several potential outcomes that might 

produce complicating aspects for nanoparticle targeting to cancers. Nanoparticle materials might be 

immediately recognized and cleared by cells of the innate immune system, limiting the usefulness 

of even their initial application. Alternately, only a fraction of the applied nanoparticles might 

engage TLR that would not significantly affect the effectiveness of the administration. Recognition 

of even a small fraction of the administered nanoparticles, however, might lead to immune events 

that diminish the effectiveness of subsequent administrations. Nanoparticles that the naïve body 

initially tolerates may become a focus of immune responses upon repeated exposure. 


26 

Oppositely, one could envisage how potential recognition by TLR proteins might be beneficial 

in certain applications. In particular, cancer cells or tumors sites that express a particular TLR might 

actually provide a unique targeting aspect for nanoparticles prepared from the right material. In 

fact, some studies have been described using empty RNA virus capsules from cowpea mosaic virus 

as biological nanoparticles for delivery. 50 Because tumors can also contain a number of immune 

cells with some that might express TLR, the potential exists to have inherent targeting of nanoparticles 

to cancers through these tumor-associated cells. Such a circumstance could reduce the 

complexity of construction for a targeted nanoparticle complex. Such a suggestion has solid 

grounding in the extensive use of nanoparticles as adjuvants used for vaccination. 51 Additionally, 

nanoparticles such as liposomes can be selectively directed specifically to Langerhans cells in the 

skin as a means of enhancing the delivery of antigens to these professional antigen presenting 

cells. 52 

Recognition of nanoparticles by TLR proteins may or may not act to stimulate potential target 

cells. If stimulation occurs, the result can include release of cytokines and a variety of potent 

regulatory molecules. Induction of pro-inflammatory cytokines may provide an undesirable 

outcome because cancerous states appear to be motivated by inflammatory events. 53 The potential 

for a beneficial or negative impact of such outcomes would require case-by-case scrutiny. By 

themselves, nanoparticles have been shown to stimulate several cell-signaling pathways, including 

those that drive the release of pro-inflammatory cytokines such as interleukin (IL)-6 and IL-8. 54,55 

The incorporation of some therapeutic or diagnostic agents into nanoparticles can further enhance 

this potential for an inflammatory outcome. 56 Additionally, a variety of nanoparticles, including 

fullerenes and quantum dots, has been shown to stimulate the creation of reactive oxygen species, 

and this effect can be enhanced by exposure to UV light that might be used for visualization and/or 

activation (reviewed in Oberdorster, Oberdorster, and Oberdorster 57 ). As discussed at the beginning 

of this section, the potential for such events may be desirable for induction of selected events, e.g., 

immunization against cancer cell antigens. Nanoparticles can also be prepared to release cytokines 

that might affect an anti-tumor immune events event. 58 

3.3.3 SAFETY ISSUES 


Although generalizations can be made for a particular targeted nanoparticle delivery system, 

specific issues will arise for each type of payload it contains and for each indication. It will be 

important to balance potential safety concerns for using nanoparticles with their potential benefits 

for reducing a safety concern that occurs without their use for comparable (or even improved) 

efficacy. Nanoparticles can tremendously reduce toxicity by sequestering cytotoxic materials from 

non-specific tissues and organs of the body until reaching the cancer site. 59 Polymeric nanoparticles 

have been loaded with tamoxifen for targeted delivery to breast cancer cells to improve the efficacy 

to safety quotient relative to direct administration of this cytotoxic agent. 60 Coupling of doxorubicin-loaded 

liposomes to antibodies or antibody fragments that can enhance targeting to cancer 

cells appears quite promising as a means of further improving the efficacy to toxicity profile for this 

chemotherapeutic. 61 PEGylated liposomal doxorubicin has improved tolerability with similar efficacy 

compared to free drug. 62 Liposomal formulations of anthracyclines appear to improve the 

cardiotoxicity profile of this anti-neoplastic. 63 PAMAM dendrimers conjugated to cisplatin not only 

improved the water solubility characteristics of this chemotherapeutic but also improve its toxicity 

profile. 64 PAMAM dendrimers have also been used to increase the effectiveness of a radioimmunotherapy 

approach by increasing the specific accumulation of radioactive atoms at a tumor site as 

well as improving selectivity of biodistribution. 65 

As previously mentioned, potential inherent targeting of some nanoparticles may initiate 

cellular responses, following recognition by immune cell surface receptor systems. Outcomes 

from such events could produce safety concerns, particularly for repeated exposures. Data 

suggest additional safety concerns related to the material(s) used in nanoparticle preparation as 



well as nanoparticle size for some routes of administration. Recent studies have called into question 

the safety of repeated aerosolized exposure of carbon nanomaterials, 66 and nanoparticles have been 

shown to have a higher inflammatory potential per given mass than do larger particles. 57 Such 

particles may provide an efficient means to actively target lung cancers that is facilitated by the 

natural properties of these materials. In general, safety issues related to using nanoparticles to target 

cancers for treatment and diagnosis will be specific for the material used to prepare the particle: its 

size, the type of targeting mechanism it utilizes, its fate at the targeted site, and its non-targeted 

distribution and elimination pattern from the body. 

3.4 TARGETING TO CANCER CELLS 

In a very simplified sense, cancers occur through dysregulation of normal cell function. The body is 

designed to correct tissue defects following damage, to expand selected immune cells in response to 

a pathogen, and to compensate for perceived deficiencies in one cell type by altering the phenotype 

of another such as in stem cell recruitment. Each of these processes is mediated by (normally) 

regulated events that allow cells to lose their differentiated phenotype with restrained growth 

characteristics and acquire a replication-driven phenotype. A lack of re-differentiation into a 

growth-restrained, differentiated phenotype is the paradigm of cancer. Repair of an epithelial 

wound is a good example of this phenomenon (Figure 3.3). It is the lack of re-differentiation 

and continued responsiveness of these cells to growth factors and growth activators that supports 

and maintains the cancer phenotype. Extensive genomic differences between differentiated and 

non-differentiated forms of the same cell account for the differences observed between these two 

cell phenotypes. 67 

(a) 

(b) 

(c) 

(d) 

Growth 

factors 

Cell division 

FIGURE 3.3 Loss and recovery of barrier function associated with normal epithelia. (a) Epithelia express tight 

junctions ( ) and associate at their base with a complex of proteins known as the basement membrane 

( ). (b) Damage to epithelia increases responses to growth factors and results in differentiated epithelial 

cells that can freely divide. (c) Cell division continues until damaged area is covered. (d) Cell–cell contacts 

such as adherens junctions, tight junctions, and gap junctions have been shown to suppress growth and 

stimulate re-differentiation. 


28 

Many of the differences between replicating and non-replicating cells are associated with 

surface and metabolic properties that can be used to discriminate between differentiated 

(normal) and de-differentiated (cancer) cells. One of the biggest concerns using this information 

to target cancer cells is that non-cancerous cells undergoing normal and necessary repair processes 

may transiently express these same targets. For example, herceptin is an antibody that binds to 

her2/neu receptors that are over expressed on the surfaces of cancer cells. Unfortunately, this 

antibody can also target normal cardiac cells undergoing repair induced by the actions of a 

common anti-cancer agent, doxorubicin; patients on doxorubicin treatment are placed on an 

extended washout period prior to exposure to herceptin. Therefore, nanoparticles having a cytotoxic 

capacity and targeted using the herceptin antibody could result in cardiomyopathy. Fortunately, the 

high, transient, systemic levels of doxorubicin associated with direct administration can be muted 

by administration in nanomaterials such as liposomes, reducing the risk of cardiomyopathy. 68 That 

many surface and metabolic properties are the same for both cancer cells of a particular cell type 

and the undifferentiated form of that cell type during normal cell function must be appreciated as 

one examines potential cancer cell-targeting strategies for nanoparticles. 

The general issues raised above concerning safety aspects of targeting cancer cells highlight 

concerns of selecting a strategy that properly accounts for unique cell surface properties and 

metabolic activities of cancer cells relative to normal cells. All too frequently, normal cells can 

undergo processes (e.g., wound repair) that will transiently transform them into a cell with surface 

properties or metabolic characteristics indistinguishable from a cancer cell. In some aspects, these 

altered surface properties and metabolic activities are intertwined. Alterations in metabolic properties 

can induce increased expression of nutrient uptake pathways and catabolic proteins that assist 

in nutrient absorption to sustain the accelerated growth rate of cancer cells. Because some of these 

components are present at the cell surface, a cancer cell’s composition and profile are modified by 

their presence. From an opposing perspective, increased surface expression of components such as 

growth factor receptors will shift the metabolic activity of a cancer cell following activation of that 

receptor (either constitutive activation or ligand-induced activation). 

3.4.1 CELL SURFACE PROPERTIES 


In general, one thinks of targeting cancer cells by use of a highly specific surface material that 

absolutely identifies only that cancer cell within the entire body. Some studies, however, have 

demonstrated that cancer cells growing in different sites of the body can have altered surface 

properties that could facilitate non-specific nanoparticle binding and uptake; colloidal iron hydroxide 

(CIH) nanoparticle association and uptake appears to be enhanced for transformed cells, 69 and 

CIH nanoparticles show differences in cell surface interactions following transformation of chick 

embryo fibroblasts. 70 Such observations suggest that generic cell-surface charge differences might 

provide a targeting strategy for nanoparticles that could be, even if not highly discriminating, used 

to enrich nanoparticle delivery to cancer cells through non-specific associations. In combination 

with other targeting strategies, surface charge differences might provide a useful adjunct. For 

example, some cancer cells express unique sets of surface enzymes that might be useful to activate 

a prodrug once a nanoparticle has been localized to the surface of a cancer cell. In this regard, a 

number of proteases have been shown to be significantly up-regulated by oncogenic conversion. 71 

A proof of concept for this type of specific application has been described using a polymer-based 

fluorogenic substrate PB-M7VIS that serves as a selective proteobeacon. 36 

A number of overexpressed growth factor receptors have been used to selectively target cancer 

cell surfaces, and the description of many of these targets has been reviewed. 72 With regard to 

targeting nanoparticles, it is important to remember that once engaged by the targeting ligand, some 

of these surface components are internalized whereas others will remain at the cell surface. 

Matching the type of nanoparticle material and its potential cargo with the likely fate of the targeted 

structure can be critical to optimizing the desired outcome. It is also possible that once bound, 



ligand–nanoparticle complexes could lead to receptor internalization that might not occur by the 

presence of the targeting ligand alone. The basis for this difference might come from the potential 

for nanoparticles to contain a coordinated ligand matrix to sequester of cell-surface receptors in a 

manner that facilitates internalization. 

Antibody-based targeting of nanoparticles to solid tumors has been a highly promising strategy 

that is augmented by the enhanced vascular permeability (EPR effect) of solid tumors. For example, 

an antibody-directed (anti-p185HER2) liposome loaded with an anti-neoplastic can be an effective 

cancer therapeutic approach. 73 Blood cell-based cancers (e.g., leukemias and lymphomas) can also 

be targeted by nanoparticles as a way to reduce unwanted systemic side effects. Nanoparticles could 

be targeted to T-cell leukemia cells using an antibody to a surface cluster of differentiation (CD) 

antigen, CD3, on the surface of lymphocytes. 74 B-cell lymphomas can be targeted by anti-idiotypic 

antibodies specific for the unique monoclonal antibody expressed by each individual cancer. 75 In 

both of these cases, the potential for these B- and T-cell-derived cancer cells to actively take up 

particles on their surfaces as part of their normal function in antigen surveillance and presentation 

might facilitate and even augment the desired outcome using a targeted nanoparticle. 

Efficient, targeted delivery of gene therapy elements and/or antigens to antigen presentation 

cells (APC) has long been a goal for the induction of anti-cancer cell immune responses. Nanoparticles 

provide an exciting possibility to achieve this goal. Coating nanoparticles with mannan 

facilitates their uptake by APCs such as macrophages and dendritic cells that acts to target these 

materials to local-draining lymph nodes following their administration. 76 Such an approach is likely 

to provide additional synergy in APC activation because polymer nanoparticles are efficiently 

phagocytosed by dendritic cells. 77 Many APCs express LDL-type receptors, and molecules that 

interact with this class of receptors could be a means of targeting as well. 78 Interestingly, LDL 

receptors can be an attractive targeting strategy for cancers because many tumors of different 

origins express elevated levels of this receptor. 79 Therefore, LDL-based nanoparticles could be 

useful in targeting cancer. 31 

It might also be possible to intentionally alter the surfaces of cancer cells to improve nanoparticle 

targeting. Cells could be transfected with a protein that expresses the appropriate acceptor 

peptide recognized by a surface-applied bacterial biotin ligase. 80 Although there would be multiple 

issues to overcome prior to clinical application, this approach outlines one strategy where cancer 

cells might be altered to express a unique surface structure such as biotin that could be very 

selectively targeted. Reversed-response targeting might also be performed using discriminating 

cell-surface properties. Hepatocytes can be targeted using nanocapsules decorated with the surface 

antigen of hepatitis B virus (SAgHBV). 81 Because liver cells may lose their capacity to bind 

SAgHBV following oncogenic conversion, this targeting strategy could be used to deliver cytoprotective 

materials to normal hepatocytes and enhance the efficacy of chemotherapeutics aimed 

at liver cancers. Similarly, hepatocytes exclusively express high affinity cell-surface receptors for 

asialoglycoproteins, and this ligand–receptor system has been used to target albumin nanoparticles 

to non-cancer cells of the liver. 8 Nanoparticles coated with galactose might also be used to target 

the liver. 82 

3.4.2 METABOLIC PROPERTIES 

One of the most detrimental aspects of cancer cells, their high rate of proliferation, can also be 

considered their Achilles’ heel. As proliferation rate increases, metabolic requirements follow 

accordingly. This places cancer cells in a precarious position where a blockade of critical metabolic 

steps can lead to cytotoxic outcomes; this is the basis from a number of currently approved anticancer 

agents that function as metabolic poisons. 83 There are two obvious approaches that could be 

used for targeting using this characteristic of cancer cells: ligands that emulate the nutrient, vitamins 

or co-factors, and antibodies that recognize these surface transport elements. Nanoparticles 

coated with a ligand for one of these receptors such as folate can be used to target to cancer cells. 84 


30 

Folate has been used to target dendrimers 85,86 and iron oxide nanoparticles. 87 Folate receptortargeted 

lipid nanoparticles for the delivery of a lipophilic paclitaxel prodrug have shown promising 

pre-clinical outcomes. 88 Cancer cells can also overexpress transferrin receptors (REF), and transferrin-conjugated 

gold nanoparticle uptake by cells has been demonstrated. 89 A transferrinmodified 

cyclodextrin polymeric nanoparticle has been described that could be used to deliver 

genetic material to cancer cells. 90 

Increased nutrient uptake associated with increased requirements for amino acids and nucleic 

acids may also provide a targeting strategy for nanoparticles. That many classical anti-cancer agents 

function by interfering with amino acid or nucleic acid incorporation into polymer structures should 

provide an important template for strategies to use nanoparticle technologies to effectively target 

cancer cells. For example, pancreatic cancer cell lines appear to overexpress the peptide transporter 

system PepT1, 91 possibly providing a growth advantage by its ability to provide additional amino 

acid uptake. Nanoparticles decorated with (or composed of) ligands recognized by this uptake 

pathway may provide an important targeting opportunity. Such an approach can be used for 

similar target-specific delivery of other molecules. It is important to remember that materials such 

as amino acids and nucleic acids, unlike co-factors discussed above that are used more as part of 

catalytic cellular events, are required in stoiciometric amounts for cell growth. Targeting strategies 

using uptake processes against vitamins and co-factors will have the added benefit of potentially 

depriving cancer cells of these critical materials whereas targeting strategies using amino acids and 

nucleic acids will probably not affect the overall influx of these materials and their incorporation into 

nascent polymers required for continued cell growth. 

A growing bank of experimental and clinical data has provided strong evidence that chronic 

inflammation can drive epithelial cell populations into an oncogenic phenotype. 92 It is this preneoplastic 

character that may act to alter the metabolic character of cells that might be useful for 

targeting nanoparticles. Such a targeting strategy would make use of transitions in cellular function 

in response to pro-inflammatory signals (Figure 3.4). In this regard, one could envisage liganddirected 

targeting to inflammatory sites as well as activation of nanoparticles (or their components) 

for localized delivery at these sites by the presence of unique enzymatic activities. Additionally, 

one could contemplate nanoparticles that deliver cancer prevention agents that work through 

suppression of inflammatory events. A ligand peptide that binds endothelial vascular adhesion 

molecule-1 (VCAM-1) on the surfaces of inflamed vessels has been used to target nanoparticles. 93 

One major concern with using such metabolic processes for targeting nanoparticles is that inflammatory 

events are a common and essential function of the body, and they are not necessarily 

associated with pre-cancerous or cancerous states. Some nanoparticles can induce an inflammatory 

response. Therefore, the method selected for such a nanoparticle-targeting strategy most keep this 

in mind. It might be possible to utilize additional cancer-targeting mechanisms (e.g., the EPR 

effect) to augment inflammation-based strategies. 

3.5 TARGETING TO TUMORS 

Nanoparticle targeting strategies involving peculiar and unique properties of tumors have been 

described. Two of these properties will be discussed: aberrant vasculature and the unique metabolic 

environment produced by inefficient blood supply resulting from the aberrant vasculature. These 

targeting strategies have strong similarities to those described for targeting cancer cells, taking 

advantage of unique endothelial cell-surface properties in tumors or the cells’ peculiar environment 

as they respond to the metabolically overactive environment of a tumor. 

3.5.1 ABERRANT VASCULATURE 


Tumor neovasculature is a promising site for targeting nanoparticles for both diagnosis and therapy. 

Once a tumor reaches a size of greater than w1 cm 3 , vascular assistance to deliver oxygen and 



(a) 

(b) 

(c) 

Macropage 

Macropage 

Cell division 

Macropage 

FIGURE 3.4 Effect of chronic pro-inflammatory stimulus on epithelial barrier patency. (a) A variety of stimuli 

( ) can incite the release of pro-inflammatory cytokine (e.g., TNFa and IFNg) from macrophages and other 

cells. Genetic predisposition can enhance these responses. (b) Released pro-inflammatory cytokines act to open 

tight junction (TJ) structures ( ), allowing entry of additional activating stimuli that, in turn, attract more cells 

associated with inflammation. (c) Chronic inflammation leads to breakdown of basement membrane, loss of TJ 

function, and disorganization of the epithelia characteristic with a pre-neoplastic state. 

remove by-products is required for cell survival. Cancer cells secrete a variety of growth factors 

that stimulate the formation of nascent blood and lymph vessels. Tumor-associated vascular beds 

have unique surface properties that are associated with their rapid growth characteristics, resulting 

in aberrant vascular beds that have been used to design a combined vascular imaging and therapy 

approach using nanoparticles. 94 Non-specific targeting of nanoparticles (10–500 nm in diameter) to 

solid tumors through this EPR capacity of solid tumors 47 can provide a means to enrich the 

localization of an anti-neoplastic agent to a tumor when it is coupled to a nanoparticle compared 

to its free form. 95 The EPR effect has also been used to increase localization through inherent 

targeting 96 of long-circulating liposomes. 97 

PAMAM dendrimers useful for boron neutron capture therapy (BNCT) of cancers have been 

targeted to tumor vasculature by attachment of vascular endothelial growth factor (VEGF) that acts 

to target VEGF receptors that are frequently overexpressed on tumor neovasculature. 98 Targeting 

VEGF receptors Flk or Flt on tumor-associated endothelial cells could also be effective; this has 

been done with a complex material composed of anti-Flk-1 antibody-coated 90 Y-labeled nanoparticles 

99 . PECAM (or CD31) is highly expressed on the surface of endothelial cells present in 

immature vasculature. Platelet endothelial cell adhesion molecule (PECAM) up-regulation 

occurs following VEGF stimulation of endothelial cells. The presence of such endothelial 

surface markers also provides the opportunity to target nanoparticles that contain DNA for gene 

therapy applications, 100 release anti-angiogenesis agents as well as chemotherapeutics, 101 and 


32 

antigens/agents to stimulate an anti-cancer cell immune response. 101,102 Vascular cell adhesion 

molecule-1 (VCAM-1) is a marker for inflammation of the endothelial and has been used to target 

nanoparticles to these sites for magento-optical imaging. 103 

Differences in tissue-specific endothelial surface properties might be used for targeting nanoparticles. 

104 In one such example, a peptide that binds to membrane dipeptidase on lung endothelial 

cells can be used to target nanocrystals to the lung. 105 Doxorubicin-loaded nanoparticles have been 

targeted to tumor vasculature by surface decoration with a cyclic arginine–glycine–aspartic acid 

(RGD) peptide that binds to the cell adhesion molecule integrin avb3 on the surface of endothelial 

cells. 106 A peptide that interacts with the lymphatic vessel marker podoplanin can be used to target 

nanocrystals to lymph vessels and some tumor cells. 105 Cationic nanoparticles, containing genes 

that can block endothelial cell signaling that were selectively targeted to tumor vasculature by 

couplingtoanintegrina vb 3 ligand, were shown to produce endothelial apoptosis and tumor 

regression. 107 

3.5.2 METABOLIC ENVIRONMENT 

Solid tumors typically grow at such a rapid pace that vascular function fails to keep pace with local 

metabolic requirements. This situation, leading to reduced oxygen tension and a depressed pH as a 

result of a lack of waste acid clearance, is associated with the onset of necrotic cores of large 

tumors. A number of studies have shown that cancer cells become adapted to successful growth in 

these difficult conditions, providing unique characteristics that might be exploited for tumor 

targeting. For example, nanoparticles will accumulate at inflammation sites along the bowel 

when administered orally. 108,109 Because these inflammatory sites are typically associated with a 

slightly acidic environment, a similar targeting strategy may be possible for tumors where the pH 

has also been depressed. Nanoparticles targeted to cancer cells might also be modified at low 

oxygen tension to either release or activate a payload. Frequently, cancer cells rely more on 

glycolysis than oxidative metabolism for energy production. This finding is consistent with the 

reduced oxygen tension of poorly perfused tumors. These conditions would result in the release of 

acidic glycolytic end-products that are not effectively cleared by the sluggish blood flow through 

tumor vascular beds. Therefore, targeting that takes advantage of a slightly depressed local pH that 

might be found in some tumors could be an attractive design strategy. 

3.6 FATE OF NANOPARTICLES 


Successful targeting of nanoparticles to cancers or tumors may involve overcoming multiple 

biological, physiological, and physical barriers. Site of initial application can be critical. For 

example, nanoparticles administered into the gut or lung would initially confront epithelial barriers. 

Metabolic events or cellular responses at an injection site represent another initial barrier to targeted 

nanoparticle delivery. Once nanoparticles have entered the body, their size, shape, or surface 

characteristics can initiate events that present a second barrier to targeted delivery—misdirection 

of the material away from its targeted site through undesired interactions. Access and/or enriched 

distribution to specific organs or regions of the body may be critical for successful nanoparticle 

targeting. Finally, once nanoparticles have reached a targeted site, metabolic or physical aspects of 

the cancer cell or tumor might limit their effectiveness. Here, the intracellular uptake and fate may 

dictate the potential success of each nanoparticle approach. Events at each of these barriers act in a 

cumulative fashion to limit the success of any nanoparticle-based targeting strategy. Obviously, the 

overall fate of nanoparticles might be improved by using materials that are not affected by these 

barriers or by modification of nanoparticles in ways that can neutralize these barrier issues. 

Matching the size and composition profile of a nanoparticle delivery system with the targeting 

strategy allows for an optimized approach that takes advantage of default targeting events as much 

as possible. 



As previously discussed, it is critical that inherent targeting mechanisms associated with a 

particular nanoparticle does not overwhelm or work in concert with any applied targeting strategy. 

Once at the target site, the fate of the nanoparticle can significantly affect its potential to provide the 

desired outcome. The delivery of hybridization-competent antisense oligonucleotides (ODNs) 

targeted to a cancer cell or tumor would not provide the desired outcome unless this material is 

efficiently internalized. ODNs covalently conjugated to anionic dendrimers have been shown to 

effectively deliver through an endocytosis process and down-regulate epidermal growth factor 

receptor expression in cancer cells. 110 Other targeting strategies may not provide the desired 

outcome if the nanoparticle is internalized by cells. For example, enzyme-coupled nanoparticles 

targeted to a tumor that would activate a prodrug could fail to provide a desired outcome. Therefore, 

the potential for nanoparticles to have a successful outcome of targeting cancer cells or tumors 

requires favorable events at natural barriers of the body and their distribution within the body, but 

also their fate at the targeted site. 

3.6.1 SITE OF INITIAL APPLICATION 

Several extracellular barriers exist for the administration and targeted delivery of nanoparticles. 

Initial entry into the body represent on obvious barrier. This is not an issue for situations where the 

cancer cell or tumor targeted is readily accessible by a topical application. Such a situation, 

however, is rather rare. Entry into the body across mucosal surfaces such as those in the gut or 

lung is typically very inefficient. Even viral particles are not very successful at this approach with 

most relying upon infecting cells of the barrier from the apical exposure by only a few viral particles 

that can replicate inside the cells to allow the basolateral (systemic) release of large numbers of 

progeny. Viral particle entry at apical surfaces of epithelial cells is decreased by physical barriers 

such as secreted mucus as well as proteases and other enzymatic barriers. Extracellular (acellular) 

matrix environments that viruses might encounter after systemic infection could similarly act to 

diminish cellular targeting and entry. Man-made nanoparticle delivery systems are likely to be 

impeded by these same physical and biological barriers at epithelial surfaces and within the body. 

Reduced surface exposure of highly charged or protruding structures is commonly used by viruses 

to minimize the impact of these extracellular barriers on viral infectivity. Similar considerations 

may facilitate optimization of nanoparticle delivery strategies. 

There are several common methods for administering materials to the body: injection or application 

to an epithelial surface (skin, intestine, lung, etc.) of the body. Nanoparticles can be absorbed 

into the skin after topical application. 111 Although nanoparticles can be taken up through appendages 

of the skin (sweat gland ducts, hair follicles) following topical application, 112 microneedles can be 

used to dramatically increase the efficiency of their uptake into and across skin. 113 The intravitreous 

injection of nanoparticles results in transretinal movement with a preferential localization in retinal 

pigment epithelial cells, 114 allowing for a sustained delivery strategy to the inner eye. 

Nanoparticles can be absorbed from the lumen of the gut, but this absorption is inefficient. 115 A 

number of factors have been examined related to regulation of nanoparticle uptake from the gut 

lumen. 116 Nanometer-sized liposomes enter into the intestinal mucosa better than larger, multilamellar 

liposomes, and this uptake can be improved by coating with a mucoadhesive polymer such 

as chitosan. 117 It is interesting that lipid-based materials absorbed from the gut partition into the 

lymphatic system and studies have suggested that these particles have remarkable access to the 

hepatocytes. 118 One way to potentially improve nanoparticle uptake from the gut is to PEGylate 

these materials in a manner that selectively increases binding to the intestinal mucosa rather than 

the stomach wall. 119 Additionally, anionic PAMAM dendrimers have been shown to rapidly cross 

the intestinal mucosa in vitro and may provide a method to improving oral delivery of nanoparticles. 

120 Cationic dendrimers also show a transcytosis capability in vitro; in general, cationic 

dendrimers are more cytotoxic than anionic dendrimers, but this characteristic can be reduced by 

additional surface modifications using lipids. 121 Formulation studies have been performed to 


34 

identify optimal methods for aerosol delivery of nanoparticles to the lung. 122 In general, the uptake 

of nanoparticles at the lung or gut surface occurs, but the efficiency of this uptake is dramatically 

improved by incorporation of a specific uptake mechanism. Even without a specific uptake 

mechanism, an appreciable amount of nanoparticle absorption can occur at these sites if they are 

sufficiently stable and are not removed by clearance mechanisms. 

A large number of studies have been performed to assess nanoparticle absorption following 

inhalation exposure, and concerns over the safety of such an approach for drug delivery have been 

raised. 123 Nanoparticles deposited in the airways appear to be taken up through transcytosis pathways 

that allow the passage of these materials across epithelial and endothelial cells to reach the blood and 

lymphatics. 57 Surface properties of nanoparticles greatly affect the capacity of nanoparticles to be taken 

into cells through the process of endocytosis and uptake following pulmonary deposition. As part of the 

respiratory tree, intranasal administration of nanoparticles can potentially provide a route into the brain. 

3.6.2 DISTRIBUTION TO SPECIFIC ORGANS 

One of the most difficult challenges of administering cytotoxic chemotherapeutics involves 

unwanted exposure to non-cancer cells and to tissue and organ compartments not involved with 

the disease. Nanoparticles carrying a chemotherapeutic can reduce the undesirable distribution of 

such compounds as they are restricted from some compartments of the body such as the brain. 59 

Oppositely, nanoparticles can be modified, e.g., conjugation to chelators, to acquire the capacity to 

transport across the blood–brain barrier or BBB. 124 Alternately, nanoparticles coupled to certain 

protein ligands such as apolipoprotein E can be used to target and transport across the BBB. 125 In 

both cases, these nanoparticle delivery approaches lead to the unique distributions of materials that 

must be cleared and/or metabolized. Therefore, one consequence of targeting cancers cells is that 

the fate of these materials, by accessing and localizing to sites where cancer cells reside, may be 

affected that could affect their overall safety as well as efficacy. 

Targeting to some cancers may require overcoming additional hurdles beyond interaction with 

specific cancer cell or tumor components. Some tumors are located in difficult-to-reach sites such as 

the brain and testes. Accessing these sites from the systemic vasculature requires that nanoparticle 

materials must first avoid systemic clearance by the RES and have the capacity to move across 

either the blood–brain or blood–testes barrier. Whereas some types of nanoparticles can keep 

materials out of compartments of the body such as the brain, 59 other types of nanoparticles may 

provide access to this difficult-to-reach compartment. 126 In general, surface characteristics that can 

be altered through chemical modifications can be used to regulate the targeted delivery of nanoparticles 

to specific sites within the body such as the brain (reviewed in Olivier 14 ). It has even been 

reported that coating nanoparticles with polysorbate 80 can facilitate brain targeting by enhancing 

their interaction with brain microvasculature. 127 Alternately, intranasal delivery of macromolecules 

has been suggested to move in a retrograde fashion into the brain following uptake at the olfactory 

epithelium. 128 Polylactic acid–PEG nanoparticles have been shown to transport across the nasal 

mucosa 129 and could, theoretically, also provide some access to the brain because both materials 

appear to transport via a transcytosis mechanism. 

3.6.3 INTRACELLULAR UPTAKE AND FATE 


Particularly in the case of cancer therapeutics, some targeted delivery may require access to intracellular 

sites such as the nucleus or mitochondria. In general, nanoparticles with diameters less than 

50 nm can easily enter most cells. Successful cellular uptake of nanoparticle systems targeted to cancer 

cells and/or tumors, however, frequently depends upon the balance of mechanisms that act to clear 

nanoparticles from the circulation and mechanisms that allow for their retention in this compartment. 

Several clearance mechanisms exist, and loss of nanoparticles from the circulation appears to be 

dominated by macrophage uptake, following complement activation or surface opsinization. 130 



Some agents such as commonly used chemotherapeutics are capable of moving efficiently across the 

plasma membrane of cells and into the cytoplasm that allows access to intracellular organelles if the 

material can be released locally to the target site. A nanoparticle-based delivery may require the release 

of a cargo either at the cell surface or after internalization of the nanoparticle by the target cell. Uptake of 

nanoparticles into cells can occur through a clathrin-coated endocytosis event, 131 through caveolae 

structures, 132 or through uptake mechanisms that do not appear to involve clathrin or caveolae. 133 

Differences in surface properties and nanoparticle size will likely dictate the predominant route of entry 

into a cell. Such uptake mechanisms into cells following nanoparticle targeting to a cancer cell or tumor 

can involve vesicular trafficking to acidified, protease, and nuclease-enriched lysosomal compartments 

within the cell. 

Unless the nanoparticle carrying a chemotherapeutic agent can release it prior to the trafficking 

of these vesicles to destructive (lysosomal) pathways or it can avoid such a pathway once inside the 

cell, the effectiveness of the absorbed material may be dramatically reduced. Also, some new 

classes of anti-cancer agents have poor membrane permeability properties and would not readily 

leave the endosome after uptake. Furthermore, exposure of these materials to lysosomal environments 

would destroy their biological activity; nucleic acid- or peptide/protein-based therapeutics 

capable of marking a cancer cell for clearance by the immune system would be examples of this 

type of approach. In these cases, proper selection of the nanoparticle composition and characteristics 

allows these materials to escape the fate of this default uptake event. There are endogenous 

properties of some materials as well as the capacity to include specific intracellular targeting agents 

that can be matched with the intracellular delivery desired for the material being targeted. 

Following specific (or non-specific) targeting of nanoparticles to a cell, a number of events act 

to traffic these material within the cell. In general, nanoparticles that have associated with a specific 

cell-surface target are internalized through an endocytosis process that produces the formation of 

intracellular vesicles, containing the nanoparticle. Based upon their physical and chemical characteristics, 

most nanoparticles reaching these endosomal vesicles are likely delivered to lysosomes 

within cells where they would be metabolized or retained. Therefore, nanoparticles would not 

readily access the cytoplasm of target cells. An ultimate fate of lysosomal structures within targeted 

cells is not necessarily a problem. Many diagnostics have already performed their function and/or 

can still be detected within this compartment. In the case of therapeutics, many of these may have 

already been released from the nanoparticle prior to its arrival at the lysosome, and/or the materials 

are stable in this hostile environment and continue to act upon the target cell from this location. 

However, there are a number of potential therapeutic compounds that will be inactivated by this 

outcome and that require additional delivery events to achieve their optimal function on the target 

cell. Oligonucleotide delivery to tumors is one such example where the therapeutic must access the 

target cell cytoplasm for its desired action. 134 

Reduction in the rate of endocytosis of nanoparticles can be achieved by coating them with 

proteins such as lactoferrin or ceruloplasmin that act to retain the material at the cell surface. 135 One 

approach to facilitate nanoparticle delivery to the cytoplasm of a target cell is to covalently attach 

the membrane penetrating TAt peptide derived from HIV-1 to the surface of these nanoparticles. 136 

Because nanoparticles of several compositions have been shown to target to the mitochondria, 

137,138 these materials may access the cell’s cytoplasm to reach this organelle. Such an 

outcome may be driven by the physico-chemical characteristics of the nanoparticle with relation to 

the unique proton and ion gradients found in mitochondria. Modification of the nanoparticle to 

affect this inherent targeting may be important. 


Nanoparticles provide a range of new opportunities to increase the targeting of currently approved 

diagnostic and therapeutic agents to cancers. Improvements in targeting can lead not only to 


36 


increased efficiency of these agents but also to increased signal-to-noise ratios for diagnostics and 

better efficacy to toxicity ratios for therapeutics. Currently, a whole new spectrum of biopharmaceuticals 

and biotechnological agents for cancer diagnosis and therapy are also being developed. 

Some of these materials require special formulation technologies to overcome drug-associated 

problems. Although nanoparticles offer improved profiles for some currently approved diagnostic 

and therapeutic agents, many biotechnology-based materials absolutely require some method of 

delivery that compensates for their poor stability or non-selective activity in a systemic setting. 

Nanoparticles offer a set of new opportunities for the development of these agents. 29 

Nanoparticles can be prepared in such a way as to have diagnostic or therapeutic agents 

integrated into them in ways that either freely releases the agents or that requires decomposition 

of the nanoparticle for the release to occur. Because of the inherent nature of small (nanometersized) 

structures, the body can identify and respond to these as foreign. Such a response can by 

suppressed by incorporation of agents that might suppress undesirable responses, or the application 

can be matched to the nanoparticle to make use of these natural responses. It is even possible to 

modify these natural responses to better match the desired clinical outcome. Specific components 

used to prepare nanoparticles can affect not only their stability in the body but also their capacity to 

be absorbed across natural barriers of the body (e.g., BBB) as well as the inherent systemic 

distribution of the nanoparticle that might compete with or complement efforts to selective 

targeting strategies. 

Without the current fanfare related to nanotechnologies, nanoparticles have been used to 

selectively target a number of organs of the body for a number of years. Nanoparticle colloids 

were shown to have contrast media properties that related to the unique surface properties of cells in 

specific organs of the body. 139 Deviations from normal function such as oncogenic transformation 

can lead to changes in a cell’s surface properties and its capacity to interact with nanomaterials. For 

example, gadolinium-based nanoparticles are taken up by hepatocytes, and by the decreased function 

and density of cancer cells in the liver tumors, a reduced level of uptake of these particles can 

be used to identify tumors using T1-weighted images obtained from MRI. 42 

Throughout this chapter, there has been frequent referral to viral infection events as a paradigm 

for cellular and intracellular targeting strategies for nanoparticles. Indeed, these materials are very 

successful models for nanoparticle targeting because they have developed mechanisms to discriminate 

between the various cells of the body (e.g., tropism for only cells of the intestinal tract) and can 

deliver labile (polynucleic acids) payloads that dramatically affect cell function. In response to 

these nanoparticle invaders, host cells have established intricate and complicated mechanisms that 

block viral infectivity and cellular actions. Such evolutionary pressures have led to the incorporation 

of intricate and novel methods by viruses to effectively combat these protective systems 

established by host cells. It is into this environment where the virus nanoparticles and host cells 

have battled back and forth for millennia that efforts to use nanoparticles to deliver agents to 

cancers for diagnosis and/or therapy must be framed. 

Finally, it is important to sound a precautionary note for the potential to over-engineer nanoparticles. 

Nanoparticles provide a platform that can potentially be used to simultaneously function 

in targeting therapeutic molecules as well a reporter and/or imaging agents. Elegant studies have 

been performed with nanoparticles modified three, four, or even five times with materials that 

promoted active targeting and/or reduced non-specific targeting as well as corrected undesirable 

properties of residence, biodistribution, and stability. Such tour-de-force efforts would be 

considered unrealistic by pharmaceutical companies for scaling to a process that would gain 

approval from regulatory agencies. Clinical success can be demonstrated for a nanoparticle 

system only if it can get to the clinic; a viable production process is critical to this development 

path. Therefore, successful applications of nanoparticles in target cancers for therapy and diagnosis 

will require designing systems where the inherent activities and distribution of nanoparticle size 

and composition allow for minimal modifications that will translate into production process steps. 



REFERENCES 

1. Kayser, O., Lemke, A., and Hernandez-Trejo, N., The impact of nanobiotechnology on the development 

of new drug delivery systems, Curr. Pharm. Biotechnol., 6(1), 3–5, 2005. 

2. Duncan, R., The dawning era of polymer therapeutics, Nat. Rev. Drug Discov., 2(5), 347–360, 2003. 

3. Douglas, S. J., Davis, S. S., and Illum, L., Nanoparticles in drug delivery, Crit. Rev. Ther. Drug 

Carrier Syst., 3(3), 233–261, 1987. 

4. Moghimi, S. M., Hunter, A. C., and Murray, J. C., Nanomedicine: Current status and future 

prospects, FASEB J., 19(3), 311–330, 2005. 

5. Zimmer, A., Antisense oligonucleotide delivery with polyhexylcyanoacrylate nanoparticles as 

carriers, Methods, 18(3), 286–295, 1999, see also p. 322. 

6. Li, H. and Qian, Z. M., Transferrin/transferrin receptor-mediated drug delivery, Med. Res. Rev., 

22(3), 225–250, 2002. 

7. Balthasar, S. et al., Preparation and characterisation of antibody modified gelatin nanoparticles as 

drug carrier system for uptake in lymphocytes, Biomaterials, 26(15), 2723–2732, 2005. 

8. Mao, S. J. et al., Uptake of albumin nanoparticle surface modified with glycyrrhizin by primary 

cultured rat hepatocytes, World J. Gastroenterol., 11(20), 3075–3079, 2005. 

9. Kickhoefer, V. A. et al., Engineering of vault nanocapsules with enzymatic and fluorescent properties, 

Proc. Natl. Acad. Sci. USA, 102(12), 4348–4352, 2005. 

10. Benns, J. M. and Kim, S. W., Tailoring new gene delivery designs for specific targets, J. Drug 

Target., 8(1), 1–12, 2000. 

11. Prabha, S. and Labhasetwar, V., Nanoparticle-mediated wild-type p53 gene delivery results in 

sustained antiproliferative activity in breast cancer cells, Mol. Pharmacol., 1(3), 211–219, 2004. 

12. Bala, I., Hariharan, S., and Kumar, M. N., PLGA nanoparticles in drug delivery: The state of the art, 

Crit. Rev. Ther. Drug Carrier Syst., 21(5), 387–422, 2004. 

13. Bejjani, R. A. et al., Nanoparticles for gene delivery to retinal pigment epithelial cells, Mol. Vis., 11, 

124–132, 2005. 

14. Olivier, J. C., Drug transport to brain with targeted nanoparticles, NeuroRx, 2(1), 108–119, 2005. 

15. Daugherty, A. L. et al., Pharmacological modulation of the tissue response to implanted polylacticco-glycolic 

acid microspheres, Eur. J. Pharm. Biopharm., 44(1637), 89–102, 1997. 

16. Lemarchand, C. et al., Novel polyester–polysaccharide nanoparticles, Pharm. Res., 20(8), 

1284–1292, 2003. 

17. Mainardes, R. M. et al., Colloidal carriers for ophthalmic drug delivery, Curr. Drug Targets, 6(3), 

363–371, 2005. 

18. Kojima, C. et al., Synthesis of polyamidoamine dendrimers having poly(ethylene glycol) grafts and 

their ability to encapsulate anti-cancer drugs, Bioconjugate Chem., 11(6), 910–917, 2000. 

19. Jain, T. K. et al., Iron oxide nanoparticles for sustained delivery of anti-cancer agents, Mol. Pharmacol., 

2(3), 194–205, 2005. 

20. He, Q. et al., Calcium phosphate nanoparticle adjuvant, Clin. Diagn. Lab. Immunol., 7(6), 899–903, 

2000. 

21. Petrovsky, N. and Aguilar, J. C., Vaccine adjuvants: Current state and future trends, Immunol. Cell 

Biol., 82(5), 488–496, 2004. 

22. Gao, X. et al., In vivo molecular and cellular imaging with quantum dots, Curr. Opin. Biotechnol., 

16(1), 63–72, 2005. 

23. Michalet, X. et al., Quantum dots for live cells, in vivo imaging, and diagnostics, Science, 307(5709), 

538–544, 2005. 

24. Kirchner, C. et al., Cytotoxicity of colloidal CdSe and CdSe/ZnS nanoparticles, Nano Lett., 5(2), 

331–338, 2005. 

25. Ruysschaert, T. et al., Liposome-based nanocapsules, IEEE Trans. Nanobioscience, 3(1), 49–55, 

2004. 

26. Shenoy, V. S., Vijay, I. K., and Murthy, R. S., Tumour targeting: Biological factors and formulation 

advances in injectable lipid nanoparticles, J. Pharm. Pharmacol., 57(4), 411–422, 2005. 

27. Moghimi, S. M. and Hunter, A. C., Capture of stealth nanoparticles by the body’s defences, Crit. Rev. 

Ther. Drug Carrier Syst., 18(6), 527–550, 2001. 


38 


28. Allen, C. et al., Controlling the physical behavior and biological performance of liposome formulations 

through use of surface grafted poly(ethylene glycol), Biosci. Rep., 22(2), 225–250, 2002. 

29. Muller, R. H. and Keck, C. M., Challenges and solutions for the delivery of biotech drugs—A review 

of drug nanocrystal technology and lipid nanoparticles, J. Biotechnol., 113(1–3), 151–170, 2004. 

30. Emanuel, N. et al., Preparation and characterization of doxorubicin-loaded sterically stabilized 

immunoliposomes, Pharm. Res., 13(3), 352–359, 1996. 

31. Versluis, A. J. et al., Stable incorporation of a lipophilic daunorubicin prodrug into apolipoprotein 

E-exposing liposomes induces uptake of prodrug via low-density lipoprotein receptor in vivo, 

J. Pharmacol. Exp. Ther., 289(1), 1–7, 1999. 

32. Nobs, L. et al., Poly(lactic acid) nanoparticles labeled with biologically active Neutravidin for active 

targeting, Eur. J. Pharm. Biopharm., 58(3), 483–490, 2004. 

33. Gref, R. et al., Surface-engineered nanoparticles for multiple ligand coupling, Biomaterials, 24(24), 

4529–4537, 2003. 

34. Konda, S. D. et al., Specific targeting of folate-dendrimer MRI contrast agents to the high affinity 

folate receptor expressed in ovarian tumor xenografts, Magma, 12(2–3), 104–113, 2001. 

35. Kobayashi, H. and Brechbiel, M. W., Dendrimer-based macromolecular MRI contrast agents: 

Characteristics and application, Mol. Imaging, 2(1), 1–10, 2003. 

36. McIntyre, J. O. et al., Development of a novel fluorogenic proteolytic beacon for in vivo detection 

and imaging of tumour-associated matrix metalloproteinase-7 activity, Biochem. J., 377(Pt 3), 

617–628, 2004. 

37. Morawski, A. M., Lanza, G. A., and Wickline, S. A., Targeted contrast agents for magnetic resonance 

imaging and ultrasound, Curr. Opin. Biotechnol., 16(1), 89–92, 2005. 

38. Voura, E. B. et al., Tracking metastatic tumor cell extravasation with quantum dot nanocrystals and 

fluorescence emission-scanning microscopy, Nat. Med., 10(9), 993–998, 2004. 

39. Marsh, J. N. et al., Improvements in the ultrasonic contrast of targeted perfluorocarbon nanoparticles 

using an acoustic transmission line model, IEEE Trans. Ultrason. Ferroelectr. Freq. Control, 49(1), 

29–38, 2002. 

40. Kiser, P. F., Wilson, G., and Needham, D., A synthetic mimic of the secretory granule for drug 

delivery, Nature, 394(6692), 459–462, 1998. 

41. Brannon-Peppas, L. and Blanchette, J. O., Nanoparticle and targeted systems for cancer therapy, 

Adv. Drug Deliv. Rev., 56(11), 1649–1659, 2004. 

42. Mintorovitch, J. and Shamsi, K., Eovist injection and Resovist injection: Two new liver-specific 

contrast agents for MRI, Oncology (Williston Park), 14(6 Suppl. 3), 37–40, 2000. 

43. Perez, R. V. et al., Selective targeting of Kupffer cells with liposomal butyrate augments portal 

venous transfusion-induced immunosuppression, Transplantation, 65(10), 1294–1298, 1998. 

44. Kim, B. K. et al., Hydrophilized poly(lactide-co-glycolide) nanospheres with poly(ethylene oxide)– 

Poly(propylene oxide)–poly(ethylene oxide) triblock copolymer, J. Microencapsul., 21(7), 697–707, 

2004. 

45. Veronese, F. M. and Pasut, G., PEGylation, successful approach to drug delivery, Drug Discov. 

Today, 10(21), 1451–1458, 2005. 

46. Cole, C. et al., Tumor-targeted, systemic delivery of therapeutic viral vectors using hitchhiking on 

antigen-specific T cells, Nat. Med., 11(10), 1073–1081, 2005. 

47. Maeda, H. et al., Tumor vascular permeability and the EPR effect in macromolecular therapeutics: A 

review, J. Control. Release, 65(1–2), 271–284, 2000. 

48. Witvrouw, M. et al., Polyanionic (i.e., polysulfonate) dendrimers can inhibit the replication of human 

immunodeficiency virus by interfering with both virus adsorption and later steps (reverse transcriptase/integrase) 

in the virus replicative cycle, Mol. Pharmacol., 58(5), 1100–1108, 2000. 

49. Moynagh, P. N., TLR signalling and activation of IRFs: Revisiting old friends from the NF-kappaB 

pathway, Trends Immunol., 26(9), 469–476, 2005. 

50. Rae, C. S. et al., Systemic trafficking of plant virus nanoparticles in mice via the oral route, Virology, 

343(2), 224–235, 2005. 

51. Aucouturier, J., Dupuis, L., and Ganne, V., Adjuvants designed for veterinary and human vaccines, 

Vaccine, 19(17–19), 2666–2672, 2001. 

52. McGuire, M. J. et al., A library-selected, Langerhans cell-targeting peptide enhances an immune 

response, DNA Cell Biol., 23(11), 742–752, 2004. 



53. Mizukami, Y. et al., Induction of interleukin-8 preserves the angiogenic response in HIF-1alphadeficient 

colon cancer cells, Nat. Med., 11(9), 992–997, 2005. 

54. Donaldson, K. and Tran, C. L., Inflammation caused by particles and fibers, Inhal. Toxicol., 14(1), 

5–27, 2002. 

55. Wottrich, R., Diabate, S., and Krug, H. F., Biological effects of ultrafine model particles in human 

macrophages and epithelial cells in mono- and co-culture, Int. J. Hyg. Environ. Health, 207(4), 

353–361, 2004. 

56. Gopalan, B. et al., Nanoparticle based systemic gene therapy for lung cancer: Molecular mechanisms 

and strategies to suppress nanoparticle-mediated inflammatory response, Technol. Cancer Res. 

Treat., 3(6), 647–657, 2004. 

57. Oberdorster, G., Oberdorster, E., and Oberdorster, J., Nanotoxicology: An emerging discipline 

evolving from studies of ultrafine particles, Environ. Health Perspect., 113(7), 823–839, 2005. 

58. Segura, S. et al., Potential of albumin nanoparticles as carriers for interferon gamma, Drug Dev. Ind. 

Pharm., 31(3), 271–280, 2005. 

59. Hamaguchi, T. et al., NK105, a paclitaxel-incorporating micellar nanoparticle formulation, can 

extend in vivo antitumour activity and reduce the neurotoxicity of paclitaxel, Br. J. Cancer, 

92(7), 1240–1246, 2005. 

60. Shenoy, D. B. and Amiji, M. M., Poly(ethylene oxide)-modified poly(epsilon-caprolactone) nanoparticles 

for targeted delivery of tamoxifen in breast cancer, Int. J. Pharm., 293(1–2), 261–270, 

2005. 

61. Park, J. W., Benz, C. C., and Martin, F. J., Future directions of liposome- and immunoliposomebased 

cancer therapeutics, Semin. Oncol., 31(6 Suppl. 13), 196–205, 2004. 

62. Rose, P. G., Pegylated liposomal doxorubicin: Optimizing the dosing schedule in ovarian cancer, 

Oncologist, 10(3), 205–214, 2005. 

63. Robert, N. J. et al., The role of the liposomal anthracyclines and other systemic therapies in the 

management of advanced breast cancer, Semin. Oncol., 31(6 Suppl. 13), 106–146, 2004. 

64. Malik, N., Evagorou, E. G., and Duncan, R., Dendrimer-platinate: A novel approach to cancer 

chemotherapy, Anti-cancer Drugs, 10(8), 767–776, 1999. 

65. Kobayashi, H. et al., Monoclonal antibody–dendrimer conjugates enable radiolabeling of antibody 

with markedly high specific activity with minimal loss of immunoreactivity, Eur. J. Nucl. Med., 

27(9), 1334–1339, 2000. 

66. Lam, C. W. et al., Pulmonary toxicity of single-wall carbon nanotubes in mice 7 and 90 days after 

intratracheal instillation, Toxicol. Sci., 77(1), 126–134, 2004. 

67. Liu, J. J. et al., Multiclass cancer classification and biomarker discovery using GA-based algorithms, 

Bioinformatics, 21(11), 2691–2697, 2005. 

68. Ewer, M. S. et al., Cardiac safety of liposomal anthracyclines, Semin. Oncol., 31(6 Suppl. 13), 

161–181, 2004. 

69. Harlos, J. P. and Weiss, L., Differences in the peripheries of Lewis lung tumor cells growing in 

different sites in the mouse, Int. J. Cancer, 32(6), 745–750, 1983. 

70. Subjeck, J. R., Weiss, L., and Warren, L., Colloidal iron hydroxide-binding to the surfaces of chick 

embryo fibroblasts transformed by wild-type and a temperature-sensitive mutant of Rous sarcoma 

virus, J. Cell. Physiol., 91(3), 329–334, 1977. 

71. Lynch, C. C. et al., MMP-7 promotes prostate cancer-induced osteolysis via the solubilization of 

RANKL, Cancer Cell, 7(5), 485–496, 2005. 

72. Carter, P., Smith, L., and Ryan, M., Identification and validation of cell surface antigens for antibody 

targeting in oncology, Endocr. Relat. Cancer, 11(4), 659–687, 2004. 

73. Park, J. W. et al., Development of anti-p185HER2 immunoliposomes for cancer therapy, Proc. Natl 

Acad. Sci. USA, 92(5), 1327–1331, 1995. 

74. Dinauer, N. et al., Selective targeting of antibody-conjugated nanoparticles to leukemic cells and 

primary T-lymphocytes, Biomaterials, 26(29), 5898–5906, 2005. 

75. Brown, S. L., Miller, R. A., and Levy, R., Antiidiotype antibody therapy of B-cell lymphoma, Semin. 

Oncol., 16(3), 199–210, 1989. 

76. Cui, Z., Han, S. J., and Huang, L., Coating of mannan on LPD particles containing HPV E7 peptide 

significantly enhances immunity against HPV-positive tumor, Pharm. Res., 21(6), 1018–1025, 2004. 


40 


77. Elamanchili, P. et al., Characterization of poly(D,L- lactic-co-glycolic acid) based nanoparticulate 

system for enhanced delivery of antigens to dendritic cells, Vaccine, 22(19), 2406–2412, 2004. 

78. Mrsny, R. J. et al., Bacterial toxins as tools for mucosal vaccination, Drug Discov. Today, 7(4), 

247–258, 2002. 

79. Vitols, S., Uptake of low-density lipoprotein by malignant cells—Possible therapeutic applications, 

Cancer Cells, 3(12), 488–495, 1991. 

80. Howarth, M. et al., Targeting quantum dots to surface proteins in living cells with biotin ligase, Proc. 

Natl Acad. Sci. USA, 102(21), 7583–7588, 2005. 

81. Yu, D. et al., The specific delivery of proteins to human liver cells by engineered bio-nanocapsules, 

FEBS J., 272(14), 3651–3660, 2005. 

82. Popielarski, S. R., Pun, S. H., and Davis, M. E., A nanoparticle-based model delivery system to guide 

the rational design of gene delivery to the liver. 1. Synthesis and characterization, Bioconjugate 

Chem., 16(5), 1063–1070, 2005. 

83. Ma, M. K. and McLeod, H. L., Lessons learned from the irinotecan metabolic pathway, Curr. Med. 

Chem., 10(1), 41–49, 2003. 

84. Rossin, R. et al., 64 Cu-labeled folate-conjugated shell cross-linked nanoparticles for tumor imaging 

and radiotherapy: Synthesis, radiolabeling, and biologic evaluation, J. Nucl. Med., 46(7), 

1210–1218, 2005. 

85. Kukowska-Latallo, J. F. et al., Nanoparticle targeting of anti-cancer drug improves therapeutic 

response in animal model of human epithelial cancer, Cancer Res., 65(12), 5317–5324, 2005. 

86. Thomas, T. P. et al., Targeting and inhibition of cell growth by an engineered dendritic nanodevice, 

J. Med. Chem., 48(11), 3729–3735, 2005. 

87. Sonvico, F. et al., Folate-conjugated iron oxide nanoparticles for solid tumor targeting as potential 

specific magnetic hyperthermia mediators: Synthesis, physicochemical characterization, and in vitro 

experiments, Bioconjugate Chem., 16(5), 1181–1188, 2005. 

88. Stevens, P. J., Sekido, M., and Lee, R. J., A folate receptor-targeted lipid nanoparticle formulation 

for a lipophilic paclitaxel prodrug, Pharm. Res., 21(12), 2153–2157, 2004. 

89. Yang, P. H. et al., Transferrin-mediated gold nanoparticle cellular uptake, Bioconjugate Chem., 

16(3), 494–496, 2005. 

90. Bellocq, N. C. et al., Transferrin-containing, cyclodextrin polymer-based particles for tumortargeted 

gene delivery, Bioconjugate Chem., 14(6), 1122–1132, 2003. 

91. Gonzalez, D. E. et al., An oligopeptide transporter is expressed at high levels in the pancreatic 

carcinoma cell lines AsPc-1 and Capan-2, Cancer Res., 58(3), 519–525, 1998. 

92. Luo, J. L., Kamata, H., and Karin, M., IKK/NF-kappaB signaling: Balancing life and death—A new 

approach to cancer therapy, J. Clin. Invest., 115(10), 2625–2632, 2005. 

93. Kelly, K. A. et al., Detection of vascular adhesion molecule-1 expression using a novel multimodal 

nanoparticle, Circ. Res., 96(3), 327–336, 2005. 

94. Li, K. C., Guccione, S., and Bednarski, M. D., Combined vascular targeted imaging and therapy: A 

paradigm for personalized treatment, J. Cell. Biochem. Suppl., 39, 65–71, 2002. 

95. Torchilin, V. P., Drug targeting, Eur. J. Pharm. Sci., 11(Suppl. 2), S81–S91, 2000. 

96. Gabizon, A. A., Selective tumor localization and improved therapeutic index of anthracyclines 

encapsulated in long-circulating liposomes, Cancer Res., 52(4), 891–896, 1992. 

97. Allen, T. M., Long-circulating (sterically stabilized) liposomes for targeted drug delivery, Trends 

Pharmacol. Sci., 15(7), 215–220, 1992. 

98. Backer, M. V. et al., Vascular endothelial growth factor selectively targets boronated dendrimers to 

tumor vasculature, Mol. Cancer Ther., 4(9), 1423–1429, 2005. 

99. Li, L. et al., A novel antiangiogenesis therapy using an integrin antagonist or anti-Flk-1 antibody 

coated 90 Y-labeled nanoparticles, Int. J. Radiat. Oncol. Biol. Phys., 58(4), 1215–1227, 2004. 

100. Konstan, M. W. et al., Compacted DNA nanoparticles administered to the nasal mucosa of cystic 

fibrosis subjects are safe and demonstrate partial to complete cystic fibrosis transmembrane regulator 

reconstitution, Hum. Gene Ther., 15(12), 1255–1269, 2004. 

101. Sengupta, S. et al., Temporal targeting of tumour cells and neovasculature with a nanoscale delivery 

system, Nature, 436(7050), 568–572, 2005. 

102. Tanaka, K. et al., Intratumoral injection of immature dendritic cells enhances antitumor effect of 

hyperthermia using magnetic nanoparticles, Int. J. Cancer, 116(4), 624–633, 2005. 



103. Tsourkas, A. et al., In vivo imaging of activated endothelium using an anti-VCAM-1 magnetooptical 

probe, Bioconjugate Chem., 16(3), 576–581, 2005. 

104. Pasqualini, R. and Ruoslahti, E., Organ targeting in vivo using phage display peptide libraries, 

Nature, 380(6572), 364–366, 1996. 

105. Akerman, M. E. et al., Nanocrystal targeting in vivo, Proc. Natl Acad. Sci. USA, 99(20), 

12617–12621, 2002. 

106. Bibby, D. C. et al., Pharmacokinetics and biodistribution of RGD-targeted doxorubicin-loaded 

nanoparticles in tumor-bearing mice, Int. J. Pharm., 293(1–2), 281–290, 2005. 

107. Hood, J. D. et al., Tumor regression by targeted gene delivery to the neovasculature, Science, 

296(5577), 2404–2407, 2002. 

108. Lamprecht, A. et al., Biodegradable nanoparticles for targeted drug delivery in treatment of inflammatory 

bowel disease, J. Pharmacol. Exp. Ther., 299(2), 775–781, 2001. 

109. Lamprecht, A. et al., A pH-sensitive microsphere system for the colon delivery of tacrolimus 

containing nanoparticles, J. Control. Release, 104(2), 337–346, 2005. 

110. Hussain, M. et al., A novel anionic dendrimer for improved cellular delivery of antisense oligonucleotides, 

J. Control. Release, 99(1), 139–155, 2004. 

111. Santos Maia, C. et al., Drug targeting by solid lipid nanoparticles for dermal use, J. Drug Target., 

10(6), 489–495, 2002. 

112. Alvarez-Roman, R. et al., Skin penetration and distribution of polymeric nanoparticles, J. Control. 

Release, 99(1), 53–62, 2004. 

113. Prausnitz, M. R. et al., Microneedles for transdermal drug delivery, Adv. Drug Deliv. Rev., 56(5), 

581–587, 2004. 

114. Bourges, J. L. et al., Ocular drug delivery targeting the retina and retinal pigment epithelium using 

polylactide nanoparticles, Invest. Ophthalmol. Vis. Sci., 44(8), 3562–3569, 2003. 

115. Florence, A. T., Issues in oral nanoparticle drug carrier uptake and targeting, J. Drug Target., 12(2), 

65–70, 2004. 

116. Florence, A. T. et al., Factors affecting the oral uptake and translocation of polystyrene nanoparticles: 

Histological and analytical evidence, J. Drug Target., 3(1), 65–70, 1995. 

117. Takeuchi, H. et al., Effectiveness of submicron-sized, chitosan-coated liposomes in oral administration 

of peptide drugs, Int. J. Pharm., 303(1–2), 160–170, 2005. 

118. Hu, J. et al., A remarkable permeability of canalicular tight junctions might facilitate retrograde, 

non-viral gene delivery to the liver via the bile duct, Gut, 54(10), 1473–1479, 2005. 

119. Yoncheva, K., Lizarraga, E., and Irache, J. M., Pegylated nanoparticles based on poly(methyl vinyl 

ether-co-maleic anhydride): Preparation and evaluation of their bioadhesive properties, Eur. 

J. Pharm. Sci., 24(5), 411–419, 2005. 

120. Wiwattanapatapee, R. et al., Anionic PAMAM dendrimers rapidly cross adult rat intestine in vitro: A 

potential oral delivery system?, Pharm. Res., 17(8), 991–998, 2000. 

121. Jevprasesphant, R. et al., Engineering of dendrimer surfaces to enhance transepithelial transport and 

reduce cytotoxicity, Pharm. Res., 20(10), 1543–1550, 2003. 

122. Sham, J. O. et al., Formulation and characterization of spray-dried powders containing nanoparticles 

for aerosol delivery to the lung, Int. J. Pharm., 269(2), 457–467, 2004. 

123. Borm, P. J. and Kreyling, W., Toxicological hazards of inhaled nanoparticles—Potential implications 

for drug delivery, J. Nanosci. Nanotechnol., 4(5), 521–531, 2004. 

124. Liu, G. et al., Nanoparticle and other metal chelation therapeutics in Alzheimer disease, Biochim. 

Biophys. Acta, 1741(3), 246–252, 2005. 

125. Muller, R. H. and Keck, C. M., Drug delivery to the brain—Realization by novel drug carriers, 

J. Nanosci. Nanotechnol., 4(5), 471–483, 2004. 

126. Garcia-Garcia, E. et al., A relevant in vitro rat model for the evaluation of blood–brain barrier 

translocation of nanoparticles, Cell. Mol. Life Sci., 62(12), 1400–1408, 2005. 

127. Sun, W. et al., Specific role of polysorbate 80 coating on the targeting of nanoparticles to the brain, 

Biomaterials, 25(15), 3065–3071, 2004. 

128. Thorne, R. G. et al., Quantitative analysis of the olfactory pathway for drug delivery to the brain, 

Brain Res., 692(1–2), 278–282, 1995. 

129. Vila, A. et al., Transport of PLA–PEG particles across the nasal mucosa: Effect of particle size and 

PEG coating density, J. Control. Release, 98(2), 231–244, 2004. 


42 


130. Moghimi, S. M. and Szebeni, J., Stealth liposomes and long circulating nanoparticles: Critical issues 

in pharmacokinetics, opsonization and protein-binding properties, Prog. Lipid Res., 42(6), 463–478, 

2003. 

131. Ma, Z. and Lim, L. Y., Uptake of chitosan and associated insulin in Caco-2 cell monolayers: 

A comparison between chitosan molecules and chitosan nanoparticles, Pharm. Res., 20(11), 

1812–1819, 2003. 

132. Gumbleton, M. et al., Targeting caveolae for vesicular drug transport, J. Control. Release, 87(1–3), 

139–151, 2003. 

133. Qaddoumi, M. G. et al., Clathrin and caveolin-1 expression in primary pigmented rabbit conjunctival 

epithelial cells: Role in PLGA nanoparticle endocytosis, Mol. Vis., 9, 559–568, 2003. 

134. Dass, C. R., Oligonucleotide delivery to tumours using macromolecular carriers, Biotechnol. Appl. 

Biochem., 40(Pt 2), 113–122, 2004. 

135. Gupta, A. K. and Curtis, A. S., Lactoferrin and ceruloplasmin derivatized superparamagnetic iron 

oxide nanoparticles for targeting cell surface receptors, Biomaterials, 25(15), 3029–3040, 2004. 

136. de la Fuente, J. M. and Berry, C. C., Tat Peptide as an efficient molecule to translocate gold 

nanoparticles into the cell nucleus, Bioconjugate Chem., 16(5), 1176–1180, 2005. 

137. Li, N. et al., Ultrafine particulate pollutants induceoxidative stress and mitochondrial damage, 

Environ. Health Perspect., 111(4), 455–460, 2003. 

138. Savic, R. et al., Micellar nanocontainers distribute to defined cytoplasmic organelles, Science, 

300(5619), 615–618, 2003. 

139. Fischer, H. W., Improvement in radiographic contrast media through the development of colloidal or 

particulate media: An analysis, J. Theor. Biol., 67(4), 653–670, 1977. 


4 

CONTENTS 

Pharmacokinetics of 

Nanocarrier-Mediated Drug 

and Gene Delivery 

Yuriko Higuchi, Shigeru Kawakami, and Mitsuru Hashida 

4.1 Introduction ...........................................................................................................................44 

4.2 Characteristics of the Structure and Pharmacokinetic Properties 

of the Nanocarrier .................................................................................................................44 

4.2.1 Linear Polymers ........................................................................................................44 

4.2.1.1 Their History and Characteristics ..............................................................44 

4.2.1.2 Pharmacokinetic Properties and Applications of Drug Carriers ...............45 

4.2.2 Polymeric Micelles....................................................................................................46 


4.2.2.2 Pharmacokinetic Properties and Applications as Drug Carriers ...............46 

4.2.3 Dendrimers ................................................................................................................47 


4.2.3.2 Pharmacokinetic Properties and Applications as Drug Carriers ...............47 

4.2.4 Liposomes..................................................................................................................48 


4.2.4.2 Pharmacokinetic Properties and Applications of Drug Carriers ...............49 

4.3 Pharmacokinetic Analysis and Tissue Distribution Characteristics 

of Macromolecules ................................................................................................................50 

4.3.1 Effect of Size on the Pharmacokinetic Properties ....................................................51 

4.3.2 Effect of Charge on the Pharmacokinetic Properties ...............................................51 

4.3.3 Effect of Modification for Cell-Selective Delivery ..................................................51 

4.4 Nanocarriers for Gene Delivery............................................................................................51 

4.4.1 Polymers ....................................................................................................................52 

4.4.2 Liposomes..................................................................................................................52 

4.4.2.1 Cationic Liposomes for Lung Targeting ...................................................52 

4.4.2.2 Mannose Liposomes for Liver NPC Targeting .........................................53 

4.4.2.3 Galactose Liposomes for Liver PC Targeting...........................................53 

4.4.2.4 Folate Receptor-Mediated Targeting .........................................................54 

References ......................................................................................................................................54 


43

44 


The recent development of nanoscale technologies is beginning to change the foundations of 

disease prevention, diagnosis, and treatment. Nanotechnology has allowed improvement of a 

novel drug carrier, nanocarrier, involving the nanoscale size and capable of targeting different 

cells in the body. These include polymeric micelles, dendrimers, and liposomes. The focus is on 

anti-tumor drugs and, although various anti-tumor drugs have been developed for cancer 

chemotherapy, such drugs often cause severe side effects related to the high cytotoxicity to 

tumor cells, and they are often toxic to normal cells. Therefore, a carefully designed nanocarrier 

is required to deliver anti-tumor drugs to their target sites for effective chemotherapy. 

Tissue accumulation and cellular uptake of externally administered agents are generally 

determined by their physicochemical and biological properties and the anatomical and physiological 

properties of the body or tissues. As far as cancer therapy is concerned, site-specific delivery of 

drugs by carrier is broadly categorized as passive and active targeting. Passive targeting is the 

method determined by the physicochemical properties of the carrier relative to the anatomical and 

physiological characteristics of the tissue. From the 1970s forward, it has been demonstrated that 

macromolecule–drug conjugates accumulate in high concentrations in solid tumors and produce a 

therapeutic effect on the cancer. 1–4 This phenomenon is dependent on the anatomical and physiological 

characteristics of the solid tumor tissue, including a large vascular permeability, high 

interstitial diffusivity, and a lack of lymphatic drainage. 5,6 As far as this effect is concerned, 

Matsumura and Maeda proposed the concept of an enhanced permeability and retention 

effect (EPR effect) of macromolecules in solid tumors. 7,8 On the other hand, active targeting 

refers to alterations in the drug carrier, using specific interactions such as ligand–receptor and 

antigen–antibody phenomena. This approach involves designing carrier materials such as lipids 

and polymers and the combination of materials that enable site-specific drug delivery at a cellular or 

sub-cellular level. 

Because the therapeutic effect of anti-tumor drugs is closely related to their pharmacokinetic 

properties, it is important for the rational design of a nanocarrier for targeted delivery to understand 

the pharmacokinetics of the nanocarrier in relation to the physicochemical and biological properties 

of the drugs involved. 9,10 In this chapter, site-specific delivery using nanocarriers is discussed based 

on pharmacokinetic considerations. As far as tissue distribution is concerned, because a nanocarrier 

possesses several features common to the macromolecule, the pharmacokinetics properties of the 

macromolecule are also summarized. 

4.2 CHARACTERISTICS OF THE STRUCTURE AND PHARMACOKINETIC 

PROPERTIES OF THE NANOCARRIER 

Recently, progress in nanotechnology has allowed the development of several types of nanocarriers 

capable of delivering drugs to target tissue and cells. Figure 4.1 summarizes typical types of 

nanocarriers and their characteristic structures. Because each type of nanocarrier has a unique 

structure and physicochemical characteristics, it exhibits unique pharmacokinetic properties in 

the body. In order to develop a strategy for establishing a nanocarrier system, it is necessary to 

understand these pharmacokinetic properties. 

4.2.1 LINEAR POLYMERS 

4.2.1.1 Their History and Characteristics 


In recent decades, the use of polymers as carriers of both covalently bound and physically entrapped 

drugs has been widely explored. The larger hydrodynamic volume of polymers contributes to the 


Pharmacokinetics of Nanocarrier-Mediated Drug and Gene Delivery 45 

Polymer conjugates 

Drug 

Dendrimers 

Polymer 

backbone 

(a) (b) 

(PAMAM dendrimers) 

N 

NH2 NH2 (c) (d) 

FIGURE 4.1 Typical nanocarriers. 

Ethylenediamine 

Methyl acrylate 

Polymeric micelles 

Liposomes 

increase in the plasma half-life of drug–polymer conjugates, increasing the probability of accumulation 

of the therapeutic agent in the tumor tissue because of the EPR effect. Drug–polymer 

conjugates (Figure 4.1) exhibit improved water solubility and reduced toxicity as a result of 

accumulation in the target tissue, and they protect the drug from enzymatic degradation or 

hydrolysis. However, drugs covalently bound to a polymer lose their pharmacological activity 

because a drug cannot interact with the target site because of steric constraints. In this respect, 

in 1975, Ringesdorf 11 was the first to propose a model for the rational design of a polymer conjugate 

with a drug that consisted of five components: the polymeric backbone, the drug, the spacer, the 

targeting group, and the solubilizing agent. The polymeric carrier can be either an inert or a 

biodegradable polymer. The drug can be fixed directly or via a spacer group onto the polymer 

backbone that is about 5–40 kDa. The proper selection of this spacer offers the possibility of 

controlling the site and the rate of release of the active drug from the conjugate by hydrolytic or 

enzymatic cleavage. 

4.2.1.2 Pharmacokinetic Properties and Applications of Drug Carriers 

Hydrophobic 

polymer 

Hydrophilic 

polymer 

Phospholipid 

Cholesterol 

To date, poly[N-(2-hydroxypropyl)methacrylamide] (PHPMA) is the most advanced example of a 

synthetic polymer used as a drug carrier in both basic research and clinical applications reported by 

Duncan 12 and Kocheck 13 . The biocompatibility of this polymer has been well characterized, 14 and 

it was found to have an inert character when it was injected in the blood stream. 15,16 A PHPMA 

copolymer conjugated with doxorubicin as an anti-tumor agent linked via a peptidyl linker 


46 

(Gly–Phe–Leu–Gly) was developed. The PHPMA copolymer conjugated with doxorubicin is 

stable in plasma 17 and has been shown to be concentrated within solid tumor models. 18,19 It is 

then cleaved intracellulary by lysosomal cysteine proteinases, 20 thereby allowing intratumoral 

drug release. Preclinical investigations have shown that a PHPMA copolymer conjugated with 

doxorubicin has radically different pharmacokinetics compared to free doxorubicin with a distribution 

plasma half-life that is increased from 5 min to 1 h. 21 The stable peptidyl linker also ensures 

that little or no free doxorubicin is released into the circulation following intravenous injection, 

increasing the therapeutic index of the conjugate. Besides the prolonged circulation in blood and 

the high solid tumor accumulation, it is also possible to attach targeting moieties such as 

antibody 22 and saccharides 23 to PHPMA. Another polymer conjugate, poly(ethylene glycol) 

(PEG), is the most basic and widely used polymer. Stylene maleic anhydride conjugated with 

the anti-tumor protein, neocarzinostatin (SMANCS), has been studied and is already used for the 

treatment of hepatocellular carcinoma. 24 Furthermore, by attaching homing devices as a cellspecific 

uptake enhancer, the biodistribution could be altered and cellular targeting of drugs 

conjugated with polymer could be achieved. 

4.2.2 POLYMERIC MICELLES 


Polymeric micelles that are prepared from amphiphilic block or graft copolymers with a spherical 

core and a shell with a carrier size of 10–100 nm have been undergoing investigation since 1984 

and have been studied more actively since 1990. 25–27 The term block refers to the linear architecture 

of the copolymer in which the end of one segment is covalently joined to the head of the other 

segment to give a diblock AB type (Figure 4.1) or multiple block (AB n) type copolymers. On the 

other hand, graft copolymers have a comb-like structure with hydrophilic segments attached on the 

side of the cationic segments. Although the influence of the copolymer architecture on biological 

activity has not yet been clarified, both block and graft copolymers can form polymeric micelles 

because of their amphiphilic character. Interactions between the polymer chains that serve as the 

driving force for micelle formation include hydrophobic, electrostatic, and p–p interactions and 

hydrogen bonding. Because hydrophobic drugs can be stably trapped in the hydrophobic core of the 

polymeric micelles and exhibit water-solubility, polymeric micelles are attractive carriers for 

hydrophobic drugs. Moreover, electrostatic (ionic) interactions involving the hydrophilic surface 

of polymeric micelles may also be applicable to macromolecules possessing many electrostatic 

(ionic) charges, e.g., DNA. 28 

4.2.2.2 Pharmacokinetic Properties and Applications as Drug Carriers 


Early reports showed that the anti-cancer drug, doxorubicin, could be incorporated into a PEGpolyaspartate 

block copolymer. 29,30 PEG constituted the outer shell of the micelle that conferred a 

stealth property on the drug. After intravenous injection, doxorubicin incorporated in polymeric 

micelles was less avidly taken up by the reticuloendothelial system (RES) and could be retained in 

the circulation for a longer time compared with an intravenous injection of doxorubicin alone. 29 

The in vivo anti-tumor activity against Colon-26 solid tumors showed that there was critical 

suppression of tumor growth and a prolonged life span for doxorubicin incorporated in polymeric 

micelles when administered to mice. 30 

Another type of polymeric micelle composed poly(ethylene oxide–aspartate)block copolymer 

and doxorubicin conjugate also exhibited a sustained circulation in blood, 31 reduced uptake by the 

RES, and more than 100 times higher accumulation in tumors in Colon-26 tumor-bearing mice. 32 

Another report described the tumor-targeting delivery of paclitaxel by block copolymers producing 

both anti-tumor activity and a reduction in a major side effect of paclitaxel, neurotoxicity. 33 



In a previous study, all-trans retinoic acid (ATRA) incorporated polymeric micelles produced a 

higher AUC and lower hepatic clearance than free ATRA, demonstrating escape from the RES. 34 

Polymeric micelles have also been used in an active targeting approach, using the fact that 

certain pathological processes are associated with local temperature increase and/or acidosis. The 

use of polymeric micelles prepared from thermo- 35,36 or pH-sensitive 37 block co-polymers allows 

the disintegration of micelles and release of the incorporated drugs specifically at the site of interest. 

Additionally, the micelles can be targeted by attaching to their surface a vector molecule such as an 

antibody, peptide, lectin, saccharide, hormone, or some low-molecular weight compounds if this 

vector molecule binds to ligands characteristic of the site of interest. 

4.2.3 DENDRIMERS 


Dendrimers are a new class of polymeric materials discovered in the 1980s that are highly 

branched, monodisperse macromolecules. These hyperbranched molecules were described by 

Tomalia et al. 38 and named dendrimers. The word dendrimer comes from the Greek dendron, 

meaning a tree or a branch, and meros meaning part. At the same time, Vogtle and his co-workers 39 

published the first report of the synthesis of polymers with a dendritic structure that they called 

cascade molecules. Newkome et al. 40 independently reported the synthesis of similar macromolecules 

that they called arboros from the Latin word arbor, meaning a tree. 

Although linear polymers are highly polydisperse with a wide range of molecular weights 

because of their automatic assembly, dendrimers are mono dispersed or low-polydispersed that 

is necessary for homogeneous and reproducible pharmacokinetic properties. Therefore, dendrimers 

that have well-defined structures and flexible functions are good candidates for drug carriers. 

Poly(amidoamine) (PAMAM) dendrimers (Figure 4.1) are the first complete dendrimer family 

to be synthesized, characterized, and commercialized. Dendrimers have a molecular architecture 

characterized by regular, dendric branching with radial symmetry. The molecular weights are 

similar to those of proteins, and their size ranges from 1 to 10 nm with each generation, or 

layer, of polymer adding 1 nm to the diameter of the molecule. 41 Lower generation dendrimers 

have a highly asymmetric shape and possess more open structures compared with those of a higher 

generation. As the chains growing from the core molecule become longer and more branched, 

dendrimers adopt a globular structure. Dendrimers become densely packed as they extend out to the 

periphery that results in a closed membrane-like structure. When a critical branched state is 

reached, dendrimers cannot continue to grow because of a lack of space. This is called the starburst 

effect. In the case of PAMAM dendrimer synthesis, this was observed after the 10th generation that 

contains 6141 monomer units and has a diameter of about 12.4 nm. 41 

4.2.3.2 Pharmacokinetic Properties and Applications as Drug Carriers 

As far as dendrimers’ pharmacokinetics are concerned, the biodistribution of [ 14 C] labeled 

PAMAM dendrimers after intravenous injection was determined for generation 3, 5, 

and 7 PAMAM dendrimers. 42 At each generation, the biodistribution characteristics were different. 

Generation 3 dendrimers showed the highest accumulation in kidney (15% of dose/g tissue over 

48 h); however, generation 5 and 7 dendrimers preferentially accumulated in the pancreas 

(peak level 32% dose/g tissue at 24 h and 20% of dose/g tissue at 2 h, respectively). In addition, 

generation 7 dendrimers showed extremely high urinary excretion with values of 46 and 74% of 

dose/g tissue at 2 and 4 h, respectively. These results suggested that the biodistribution of a 

dendrimer depended on the generation because physicochemical properties were determined by 

the generation. Other reports described the biodistribution of [ 3 H]-labeled neutral and positively 


48 

charged generation 5 PAMAM dendrimers in a tumor-bearing mouse model. 43 Both neutral and 

positively charged dendrimers showed a similar biodistribution trend 1 h after intravenous injection; 

the highest level was found in the lungs, kidney, and liver (positive, 28–6.1% of dose/g 

tissue; neutral, 5.3–2.9% of dose/g tissue) followed by the tumor, spleen, and pancreas (positive, 

3.3–2.6% of dose/g, tissue; neutral 3.4–0.92% of dose/g tissue). Differences between neutral and 

positively charged dendrimers were observed for accumulation in the lung. A high lung accumulation 

of cationic dendrimer was found compared with conventional cationic liposomes, and this 

may be explained by the stacking to the capillary vessels of the lung. Moreover, [ 14 C]-labeling 

of the dendrimer results in partial neutralization of the dendrimer. Therefore, different biodistribution 

properties were observed in the lung and liver accumulation between [ 14 C]-and [ 3 H]labeled 

dendrimers. These results suggest that the biodistribution properties of radio-labeled 

dendrimers have to be examined carefully, taking into consideration the surface charge 

caused by the radio-labeling. As far as tumor accumulation is concerned, this may be explained 

by the EPR effect. 

Drug delivery applications of dendrimers have been studied by Szoka and co-workers. 44 The 

biodistribution of doxorubicin covalently bound via a hydrazone linkage with a polyester dendritic 

scaffold based on 2,2-bis(hydroxymethyl)propanoic acid was examined. After intravenous injection, 

doxorubicin was extracted from the tumor and organs. Compared with free doxorubicin, the 

serum half-life was significantly increased. Therefore, this dendrimer would be a promising drug 

carrier for cancer therapy. 




Liposomes that are vesicles of lipid bilayers were first prepared in the 1960s. Liposome encapusulation 

has been proven useful for reducing the toxicity of drugs and increasing the solubility of 

hydrophobic drugs. However, liposomes tend to be trapped by the RES after intravenous injection. 

Therefore, allowing escape from the RES is one of the key strategies for developing cell-specific 

carriers because liposomes escaping from the RES accumulate in tumor tissue by a passive 

mechanism. As far as methods of escaping from the RES are concerned, Allen et al. in 1987 45 

were the first to show that ganglioside and sphingomyelin (SM) synergistically acted to dramatically 

reduce the rate and extent of uptake of liposomes by the RES; therefore, this type of liposome 

is called a stealth liposome. In 1990, Klibanov et al. 46 demonstrated that PEG-liposomes prepared 

as poly (ethylene glycol)–phosphatidylethanolamine conjugates (PEG–PE) could also reduce 

uptake by the RES. Because PEG is easy to prepare and relatively cheap, and it has controllable 

molecular weight and is able to bind to lipids or proteins, it can be used for active targeting such as 

glycosylated liposomes 47 and immuno-liposomes. 48 

On the other hand, as far as active targeting is concerned, the receptor-mediated endocytosis 

systems present in various cell types would be useful, and a number of gene delivery systems were 

developed in the 1990s. Ligand–receptor recognition is an attractive tool for the development of 

cell-specific targeting systems; i.e., galactose to asialoglycoprotein receptors expressed on hepatocytes 

parenchymal cell (PC), 49 mannose to mannose receptors expressed on macrophages, 

sinusoidal endothelial cells, Kupffer cells, and dendritic cells, 50 folate receptors expressed on 

cancer cells, 51 and so on. As far as the size effect is concerned, it is possible to prepare liposomes 

of the size required by filtration. Because capillary vessels in a human tumor inoculated into SCID 

mice are permeable even to liposomes up to 400 nm in diameter, 52 liposomes with a diameter less 

than 50–200 nm are typically used. On the other hand, with regard to the effect of the surface charge 

of liposomes, it is known that strongly cationic liposomes highly accumulate in the lung after 

intravenous injection. Table 4.1 summarizes liposomes for drug delivery. 



TABLE 4.1 

In Vivo Targeted Drug Delivery 

System Method Target Tissue (Cell) Reference 

EPR (prolongation of circulating) Sphingomyelin modification Tumor 45 

53 

55 

56 

PEG modification Tumor 46 

57 

58 

59 

Asialoglycoprotein receptor mediated Galactose modification Liver (hepatocyte) 62 

63 

64 

61 

65 

Mannose receptor mediated Mannose modification Liver (non-parenchymal cell) 66 

67 

71 

4.2.4.2 Pharmacokinetic Properties and Applications of Drug Carriers 

SM-liposomes described by Allen et al. dramatically reduced the uptake by the RES and exhibited a 

prolonged circulation in blood and, consequently, accumulated in tumor tissue. 45 Gabizon and 

Papahadjopoulos 53 have examined a number of other natural and synthetic lipids to prolong the 

circulation time and lead to preferential uptake by tumor cells in vivo. In these respects, SM-containing 

liposomes would be an attractive option as a tumor targeting carrier for highly lipophilic 

anti-tumor drugs. The preparation of SM containing emulsion formulations has been also described 

to prolong the blood retention of encapsulated [ 3 H] labeled ATRA. 54 Furthermore, after the intravenous 

injection of vincristine encapsulated in SM-liposomes, the plasma vincristine level was 

7-fold higher than that following administration in the form of bare liposomes, and the half-life of 

the elimination from the circulation was prolonged. 55,56 The improved circulation lifetime of 

vincristine in SM-liposomes correlated with the increased vincristine accumulation in subcutaneous 

solid A431 human xenograft tumors. Moreover, treatment with vincristine in SM-liposomes 

delayed the increase in tumor mass. 55 

PEG-liposomes are the most widely studied stealth liposomes. This occurs because the presence 

of PEG protects liposomes from the interaction with opsonins in the blood plasma and prevents their 

rapid uptake by the RES. 46,57,58 Therefore, doxorubicin encapsulated into PEG liposomes exhibited 

a dramatically increase in circulating time in blood and much better therapeutic efficacy as compared 

with the free drug. Therefore, they are currently being used in the clinic as a component of Doxil w . 59 

On the other hand, as far as active targeting is concerned, galactose modification could be an 

attractive tool for PC targeting of lipophilic drugs. Recently, a novel galactosylated cholesterol 

derivative, Gal-C4-Chol, that could be stably incorporated into liposomes by means of its lipophilic 

anchor and allowed the introduction of galactose residues on the surface of the liposomes was 

synthesized. 60,61 After intravenous injection, galactosylated liposomes prepared using Gal-C4-Chol 

were preferentially taken up by the liver. 62 In addition, prostaglandin E 63 and probucol 64 were 

investigated as model lipophilic drugs incorporated in galactosylated liposomes prepared using 

Gal-C4-Chol, and they were found to be effectively taken up by PC. Recently, other types of 

galactosylated liposome were developed. 65 


50 

Using a mannosylated cholesterol derivative (Man-C4-Chol) that was synthesized by the same 

method as Gal-C4-Chol, mannosylated liposomes (Man-liposomes) for non-parenchymal cell 

(NPC) and macrophages targeting were prepared. 66–68 As far as cancer immunotherapy was 

concerned, macrophages, Kupffer cells, and dendric cells are attractive targets. These cells 

expressed a large number of mannose receptors on the cell surface and, therefore, Man-liposomes 

would be ideal carriers for cancer therapy. Previous studies have shown that Kupffer cells 

(contained in NPC) can be activated to a tumor reducible state by the administration of immunomodulators 

such as muramyl dipeptide. 69 However, muramyl dipeptide parentally administrated in 

free form is rapidly cleared from the body and excreted in the urine. 70 It was demonstrated that 

active targeting of muramyl dipeptide to liver non-parenchymal cells by Man-liposomes resulted in 

more effective inhibition of liver metastasis than delivery of muramyl dipeptide by liposomes 

without mannose. 71 Moreover, treatment with muramyl dipeptide incorporated in Man-liposomes 

increased the survival of tumor-bearing mice. 71 

4.3 PHARMACOKINETIC ANALYSIS AND TISSUE DISTRIBUTION 

CHARACTERISTICS OF MACROMOLECULES 

The pharmacokinetic characteristics are generally determined by the physicochemical properties of 

the molecule such as the size, molecular weight, and surface charge. 3 Moreover, specific 

interactions such as ligand–receptor or antibody–antigen have a major effect on the characteristics 

of the molecule after intravenous injection. 4 The former factors are closely related to passive 

targeting, and the latter factors are closely related to active targeting. Figure 4.2 summarizes the 

Apparent hepatic uptake clearance (ml/hr) 

100 

10 

1 

0.1 

0.01 

pDNA 

PEG-CAT 

BSA 

(67,000) 

PEG-SOD 

Hepatic plasma flow rate 

Cation: 

Electrostatic interaction 

Cat-BSA 

Weak anion: 

Long circulation 

Man-SOD 

Suc-BSA 

Strong anion: 

Electrostatic interaction 

SOD 

(32,000) 

0.01 0.1 1 

10 

Urinary excretion clearance (ml/h) 

Gal-SOD 

ODN 

Inurin 

(5,200) 

Rate of fluid-phase endocytosis of liver 


Glomerular filtration rate 

Small macromolecule: 

Urinary excretion 

FIGURE 4.2 Effect of physicochemical characteristics of macromolecules on their hepatic uptake and urinary 

excretion in mice. The number in parenthesis is the molecular weight. Abbreviations in this figure are as 

follows: SOD, superoxide dismutase; Man-SOD, mannosylated superoxide dismutase; Gal-SOD, galactosylated 

superoxide dismutase; PEG-SOD, pegylated superoxide dismutase; Cat-BSA, cationized bovine serum 

albumin; PEG-CAT, pegylated catalase; ODN, oligonucleotide. (From Kawabata, K., Takakura, Y., and 

Hashida, M., Pharm. Res., 12, 1995; Hashida, M., et al., J. Control. Release, 41,1996; Fujita, T., et al., J. 

Pharmacol. Exp. Theor., 263, 1992; Yamasaki, Y., et al., J. Pharmacol. Exp. Theor., 301, 2002; Yabe, Y., 

et al., J. Pharmacol. Exp. Theor., 289, 1999.) 



effect of size, charge, and ligand modification of macromolecules on hepatic uptake clearance and 

urinary excretion clearance that essentially decide the pharmacokinetic properties of the macromolecule. 

Understanding the effect of the physicochemical properties of macromolecules on their 

tissue distribution allows theoretical design delivery systems, involving nanocarriers. 

4.3.1 EFFECT OF SIZE ON THE PHARMACOKINETIC PROPERTIES 

After intravenous injection, the kidney plays an important role in the disposition of macromolecules 

circulating in the blood. Macromolecules with a molecular weight of less than 50,000 (approximately 

6 nm in diameter) are susceptible to glomerular filtration and are excreted in the urine. 72–75 

Therefore, low molecular weight drugs and oligonucleotides with a lower molecular weight 

undergo rapid glomerular filtration by the kidney. 76,77 On the other hand, pDNA is too large to 

be filtered without degradation. 78 

4.3.2 EFFECT OF CHARGE ON THE PHARMACOKINETIC PROPERTIES 

Electric charge is also an important factor in determining the pharmacokinetic properties of macromolecules. 

79 Because the glomerular capillary walls also function as a charge-selective barrier 

having negative charges, positively charged macromolecules exhibit higher glomerular permeability 

than anionic macromolecules of similar molecular weights. Larger molecular weight 

cationic macromolecules escaping glomerular filtration mainly accumulate in the liver, kidney, 

and lung. 80–82 This phenomenon can be explained by following factors: 

1. Because the cell surface is negatively charged in general, cationic molecules tend to 

electrostatically bind to it. 

2. Because the liver and kidney have fenestrated capillaries, macromolecules can cross the 

endothelial barrier. 

3. Direct interaction with the lung endothelial cells. 

4. Embolization of aggregates of the complex with negatively charged blood components. 

On the other hand, strongly negatively charged macromolecules are taken up by liver nonparenchymal 

cells through liver fenestrae by scavenger receptor-mediated endocytosis. 78,83 Neutral 

or weakly anionic macromolecules with a molecular size that does not allow glomerular filtration 

such as polyethylene glycol and serum albumin posses a very weak affinity for the cell surface, and 

they are cleared very slowly. 7,8,84,85 Consequently, such neutral or weakly anionic macromolecules 

have a longer elimination half-life compared with the approximately same sized macromolecule 

with a strong negative or positive charge. 

4.3.3 EFFECT OF MODIFICATION FOR CELL-SELECTIVE DELIVERY 

Ligand modification for active targeting using receptor-mediated uptake markedly affects the 

pharmacokinetic properties of macromolecules. 4 For example, pharmacokinetic studies of 

receptor-mediated targeting to liver parenchymal cells via galactose recognition have been 

carried out based on the clearance concept. Galactose modification increased the hepatic uptake 

of macromolecules of different molecular weights. 3 The clearance was almost the same as the 

hepatic flow rate in mice, showing that it is almost identical to the theoretical maximum value. 

4.4 NANOCARRIERS FOR GENE DELIVERY 

In addition to traditional drugs, pDNA has recently also been used as a drug for cancer therapy; for 

example, antigen-encoding pDNA has a potential application in DNA vaccine therapy for cancer. 

Moreover, new techniques to inhibit target gene expression through transcriptional regulation 


52 

without any changes in the functions of other genes using synthetic oligonucleotides such as 

antisense DNA, decoy oligonucleotide, and siRNA have been also considered as a novel anti-tumor 

chemotherapy. However, because pDNA and synthetic oligonucleotides are easily degraded or 

metabolized, there is no cell selectivity, and enough transgene expression or inhibition of target 

gene expression for effective therapy could not be achieved by the administration of naked pDNA 

or oligonucleotides. Therefore, for effective therapy, it is essential to develop a novel carrier that 

makes it possible to deliver such drugs to the target tissue or cells. 

4.4.1 POLYMERS 

Polymeric micelles prepared from cationic copolymer could be a novel gene carrier. After 

preparing a pDNA complex (standard size 5,000–7,000 bp), the mean particle of the complex is 

about 50–200 nm. pDNA entrapped in polymeric micelles is highly resistant to nuclease 

degradation. 28 

Dendrimers are also an attractive carrier for gene delivery because they can interact on an 

electrostatic charge basis with biologically relevant polyanions such as nucleic acids 86 and pDNA 87 

because their surfaces are covered with primary amino groups. Moreover, chemical or biological 

modification of the surface of dendrimers would make them more effective drug or gene carriers. 


Liposomes with cationic charge on their surface are novel carriers for genes. Several kinds of 

liposomes for gene delivery have been developed depending on passive or active targeting 

mechanisms. Table 4.2 summarizes liposomes for gene delivery. 

4.4.2.1 Cationic Liposomes for Lung Targeting 

After intravenous administration of pDNA complex, the immediate effect on the complex of 

erythrocytes is to induce aggregation. 88 Large aggregates (O400 nm) may be readily entrapped 

in the lung capillary. Mahato et al. demonstrated that cationic liposome/[ 32 P] pDNA complexes 

(lipoplexes) were rapidly cleared from the blood circulation and accumulated extensively in the 

lung and liver. 89 In addition, [ 32 P] lipoplexes were predominantly taken up by the liver NPC and 

TABLE 4.2 

In Vivo Targeted Gene Delivery 


System Liposome Target Tissue (Cell) Reference 

Capillary trapping Cationic-liposome Lung 90 

89 

91 

93 

Asialoglycoprotein receptor madiated Asialofetuin-liposome Liver (hepatocyte) 96 

Galactose-liposome Liver (hepatocyte) (hepatocyte) 97 

60 

98 

Mannose receptor mediated Mannose-liposome Liver (non-parenchymal cell) 66 

94 

95 

Folate receptor mediated Folate-liposome Tumor cell 99 

100 



this uptake was inhibited by the pre-administration of dextran sulfate, suggesting the involvement 

of a phagocytic process. However, gene expression after intravenous administration of lipoplexes 

was extremely high in the lung. 90–92 It is well known that IFNb exhibits anti-tumor activity. Taking 

this into consideration, intravenously injected pDNA encoding IFNb complexed with cationic 

liposome was found to significantly inhibit established metastatic lung tumors in CT-26 cells 

inoculated into mice. 93 

4.4.2.2 Mannose Liposomes for Liver NPC Targeting 

Because Man-C4-Chol has a cationic charge, a cationic charge could also be induced on the 

surface of a Man-liposome. Therefore, Man-liposomes are also an attractive potential gene carrier 

because of this cationic charge that would allow electrical interaction with the gene. As far as 

in vivo gene transfection is concerned, the highest gene expression was observed in the liver after 

intravenous injection of pDNA/Man-liposome complexes (Man-lipoplex) in mice. 66–68 In 

addition, gene transfer by a Man-liposome was mannose receptor-mediated because the gene 

expression with Man-lipoplex in the liver was significantly reduced by pre-dosing with mannosylated 

bovine serum albumin. Because macrophages in the liver and spleen exist at the 

endothelial cells intervals and could make contact with the complex without passing through 

the sinusoid (100–200 nm), cell-selective gene transfection could be achieved by intravenous 

administration of Man-lipoplex. Therefore, a mannosylated gene carrier would be effective for 

the NPC-selective gene transfection system even after intravenous administration. 94 As far as the 

application to cancer therapy using targeted gene delivery by mannosylated liposomes is 

concerned, DNA vaccine therapy is suitable because antigen encoded pDNA can be efficiently 

transfected into dendritic cells that expressed a large number of mannose receptors. Recently, 

Hattori et al. 95 demonstrated that the targeted delivery of a DNA vaccine by Man-liposomes is a 

potent method for DNA vaccine therapy. 

4.4.2.3 Galactose Liposomes for Liver PC Targeting 

PC exclusively expresses large numbers of high affinity cell-surface receptors that can bind asialoglycoprotein 

and internalize to the cell interior. In order to achieve PC-specific gene transfection, 

galactose moieties are introduced into the cationic liposomes. The galactosylation of liposomes can 

be achieved by coating with either glycoproteins or galactose conjugated synthetic lipids. Hara et 

al. 96 reported that asialofetuin-labeled liposomes encapsulating pDNA were taken up by asialoglycoprotein 

receptor-mediated endocytosis using cultured PC and showed the highest hepatic gene 

expression after intraportal injection with a preloading of EDTA. Gal-C4-Chol that possess a 

similar structure to Man-C4-Chol was synthesized. 58 In vivo gene transfer was examined by 

optimizing the pharmacokinetics and physicochemical properties. 97 The radioactivity in the liver 

from the Gal-liposome/[ 32 P] pDNA complex (Gal-lipoplexes) was about 75% of the dose even 

1 min after intraportal administration. The hepatic gene expression of Gal-lipoplexes was more than 

a 10-fold greater than that of lipoplexes with bare cationic liposomes. When gene expression was 

examined at the intrahepatic cellular level, the gene expression of PC of Gal-lipoplexes was 

significantly higher than that of liver NPC. On the other hand, the gene expression of PC and 

NPC of lipoplexes was almost identical. In addition, asialoglycoprotein receptor-mediated endocytosis 

was also confirmed by the inhibitory effect of pre-dosing an excess amount of galactosylated 

bovine serum albumin. It was previously reported that the lipoplexes interact with erythrocytes after 

intravenous administration. 88 Recently, the presence of an essential amount of sodium chloride 

(NaCl) during the formation of Gal-lipoplexes stabilizing the complexes in accordance with the 

surface charge regulation (SCR) theory was demonstrated. 98 The transfection activity in hepatocytes 

of SCR Gal-lipoplexes was significantly higher than that of conventional complexes. 


54 

4.4.2.4 Folate Receptor-Mediated Targeting 

The folate receptor is known to be over expressed in a large fraction of human tumors, but it is only 

minimally distributed in normal tissues. Therefore, the folate receptor has also been used as a tumortargeting 

ligand for several drug delivery systems. Recently, Hofland et al. synthesized folate-PEGlipid 

derivatives to prepare folate-modified cationic liposomes. 99 After an intravenous injection of 

folate–conjugated liposome complex, gene expression in the tumors was not changed, whereas gene 

expression in the lung was reduced as compared with the conventional complex. However, after 

intraperitoneal injection in a murine disseminated peritoneal tumor model, folate–conjugated liposome 

complex formulations produced approximately a 10-fold increase in tumor-associated gene 

expression compared with conventional complex. 100 

REFERENCES 


1. Sezaki, H., Takakura, Y., and Hashida, M., Soluble macromolecular carriers for the delivery of 

antitumour drugs, Adv. Drug Deliv. Rev., 3, 247, 1989. 

2. Sezaki, H. and Hashida, M., Macromolecule–drug conjugates in targeted cancer chemotherapy, Crit. 

Rev. Theor. Drug Carrier Syst., 1, 1, 1984. 

3. Takakura, Y. and Hashida, M., Macromolecular carrier systems for targeted drug delivery: Pharmacokinetic 

considerations on biodistribution, Pharm. Res., 13, 820, 1996. 

4. Hashida, M., Kawakami, S., and Yamashita, F., Lipid carrier systems for targeted drug and gene 

delivery, Chem. Pharm. Bull., 53, 871, 2005. 

5. Takakura, Y. and Hashida, M., Macromolecular drug carrier systems in cancer chemotherapy: 

Macromolecular prodrugs, Crit. Rev. Oncol. Hematol., 18, 207, 1995. 

6. Jain, R. K., Transport of molecules across tumor vasculature, Cancer Metastasis Rev., 6, 559, 1987. 

7. Matsumura, Y. and Maeda, H., A new concept for macromolecular therapeutics in cancer 

chemotherapy: Mechanism of tumoritropic accumulation of proteins and the anti-tumor agent 

smancs, Cancer Res., 46, 6387, 1986. 

8. Maeda, H. and Matsumura, Y., Tumoritropic and lymphotropic principles of macromolecular drugs, 

Crit. Rev. Theor. Drug Carrier Syst., 6, 193, 1989. 

9. Takakura, Y. et al., Influence of physicochemical properties on pharmacokinetics of non-viral 

vectors for gene delivery, J. Drug Target., 10, 99, 2002. 

10. Nishikawa, M. and Hashida, M., Pharmacokinetics of anti-cancer drugs, plasmid DNA, and their 

delivery systems in tissue-isolated perfused tumors, Adv. Drug Deliv. Rev., 40, 19, 1999. 

11. Ringsdorf, H., Structure and properties of pharmacologically active polymers, J. Polym. Sci. Polym. 

Symp., 51, 135, 1975. 

12. Duncan, R. et al., Polymers containing enzymatically degradable bonds 7. Design of oligopeptide 

sidechains in poly(N-(2-hydroxypropyl) metacylamide) copolymers to promote efficient degradation 

by lysosomal enzymes, Makromol. Chem., 184, 1997–2008, 1983. 

13. Kopecek, J. and Bazilova, H., Poly[N-(hydroxypropyl)-methacrylamide]—1. Radical polymerisation 

and copolymerization, Eur. Polym. J., 9, 7, 1973. 

14. Rihova, B. et al., Biocompatibility of N-(2-hydroxypropyl) methacrylamide copolymers containing 

adriamycin, Biomaterials, 10, 335, 1989. 

15. Cartlidge, S. A. et al., Soluble, crosslinked N-(2-hydroxypropyl) methacrylamide copolymers as 

potential drug carriers 2: Effect of molecular weight on blood clearance and body distribution in 

the rat after intravenous administration. Distribution of unfractionated copolymer after intraperitoneal, 

subcutaneous or oral administration, J. Control Release, 4, 253, 1987. 

16. Seymour, L. W. et al., Effect of molecular weight of N-(2-hydroxypropyl)methacrylamide copolymers 

on body distribution and rate of excretion after subcutaneous, intraperitoneal, and intravenous 

administration to rats, J. Biomed. Mater. Res., 21, 1341, 1987. 

17. Rejmanova, P. et al., Stability in rat plasma and serum of lysosomally degradable oligopeptide 

sequences in N-(2-hydroxypropyl)methacrylamide) copolymers, Biomaterials, 6, 45, 1985. 

18. Seymour, L. W. et al., Tumour tropism and anti-cancer efficacy of polymer-based doxorubicin 

prodrugs in the treatment of subcutaneous murine B16F10 melanoma, Br. J. Cancer, 70, 636, 1994. 



19. Cassidy, J. et al., Clinical trials of nimodipine as a potential neuroprotector in ovarian cancer patients 

treated with cisplatin, Cancer Chemother. Pharmacol., 44, 161, 1998. 

20. Duncan, R. et al., Degradation of side-chains of N-(2-hydroxypropyl)methacrylamide) copolymers 

by lysosomal thiolproteinases, Biosci. Rep., 2, 1041, 1982. 

21. Seymour, L. W. et al., Pharmacokinetics of polymer-bound Adriamycin, Biochem. Pharmacol., 39, 

1125, 1990. 

22. Flanagan, P. A. et al., Evaluation of protein-N-(2-hydroxypropyl) methacrylamide copolymer conjugates 

as targetable drug carriers 2. Body distribution of conjugates containing transferrin, antitransferrin 

receptor antibody or anti-Thy 1.2 antibody and effectiveness of transferrin-containing 

daunomycin conjugates against mouse L 1210 leukaemia in vivo, J. Control Release, 18, 25, 1992. 

23. Wedge, S. R., Duncan, R., and Kopeckova, P., Comparison of the liver subcelluar distribution of free 

daunomycin and that bound to galactosamine targeted N-(2-hydroxypropyl)methacrylamide copolymers, 

following intravenous administration in the rat, Br. J. Cancer, 63, 546, 1991. 

24. Maeda, H., Sawa, T., and Konno, T., Mechanism of tumor-targeted delivery of macromolecular 

drugs, including the EPR effect in solid tumor and clinical overview of the prototype polymeric drug 

SMANCS, J. Control Release, 74, 47, 2001. 

25. Yokoyama, M., Block copolymers as drug carriers, Crit. Rev. Theor. Drug Carrier Syst., 9, 213, 

1992. 

26. Kataoka, K. et al., Block copolymer micelles as vehicles for drug delivery, J. Control Release, 24, 

119, 1993. 

27. Torchilin, V. P., Structure and design of polymeric surfactant-based drug delivery systems, 

J. Control Release, 73, 137, 2001. 

28. Kakizawa, Y. and Kataoka, K., Block copolymer micelles for delivery of gene and related 

compounds, Adv. Drug Deliv. Rev., 54, 203, 2002. 

29. Yokoyama, M. et al., Selective delivery of adriamycin to a solid tumor using a polymeric micelle 

carrier system, J. Drug Target., 7, 171, 1999. 

30. Yokoyama, M. et al., Toxicity and anti-tumor activity against solid tumors of micelle-forming 

polymeric anti-cancer drug and its extremely long circulation in blood, Cancer Res., 51, 3229, 1991. 

31. Kwon, G. et al., Biodistribution of micelle-forming polymer–drug conjugates, Pharm. Res., 10, 970, 

1993. 

32. Kwon, G. et al., Enhanced tumor accumulation and prolonged circulation times of micelle-forming 

poly (ethylene oxide–aspartate) block copolymer–adriamycin conjugates, J. Control Release, 29, 17, 

1994. 

33. Hamaguchi, T. et al., NK105, a paclitaxel-incorporating micellar nanoparticle formulation, can 

extend in vivo antitumour activity and reduce the neurotoxicity of paclitaxel, Br. J. Cancer, 92, 

1240, 2005. 

34. Kawakami, S. et al., Biodistribution characteristics of all-trans retinoic acid incorporated in liposomes 

and polymeric micelles following intravenous administration, J. Pharm. Sci., 94, 2606, 2005. 

35. Kohori, F. et al., Preparation and characterization of thermally responsive block copolymer micelles 

comprising poly(N-isopropylacrylamide-b-DL-lactide), J. Control Release, 55, 87, 1998. 

36. Chung, J. E. et al., Effect of molecular architecture of hydrophobically modified poly(N-isopropylacrylamide) 

on the formation of thermoresponsive core-shell micellar drug carriers, J. Control 

Release, 53, 119, 1998. 

37. Le Garrec, D. et al., Optimizing pH-responsive polymeric micelles for drug delivery in a cancer 

photodynamic therapy model, J. Drug Target., 10, 429, 2002. 

38. Tomalia, D. A. et al., A new class of polymers: Starburst-dendritic macromolecules, Polym. J., 17, 

117, 1985. 

39. Buhleier, E., Wehner, W., and Vogtle, F., Synthesis, 405, 155, 1978. 

40. Newkome, G. R. et al., Cascade molecules: A new approach to micelles, A[27]-arborol, J. Org. 

Chem., 50, 2003–2004, 1985. 

41. Svenson, S. and Tomalia, D. A., Dendrimers in biomedical applications-reflections on the field, Adv. 

Drug Deliv. Rev., 57, 2106, 2005. 

42. Roberts, J. C., Bhalgat, M. K., and Zera, R. T., Preliminary biological evaluation of polyamidoamine 

(PAMAM) Starburst dendrimers, J. Biomed. Mater. Res., 30, 53, 1996. 


56 


43. Nigavekar, S. S. et al., 3 H dendrimer nanoparticle organ/tumor distribution, Pharm. Res., 21, 476, 

2004. 

44. Padilla De Jesus, O. L. et al., Polyester dendritic systems for drug delivery applications: In vitro and 

in vivo evaluation, Bioconjug. Chem., 13, 453, 2002. 

45. Allen, T. M. and Chonn, A., Large unilamellar liposomes with low uptake into the reticuloendothelial 

system, FEBS Lett., 223, 42, 1987. 

46. Klibanov, A. L. et al., Amphipathic polyethyleneglycols effectively prolong the circulation time of 

liposomes, FEBS Lett., 268, 235, 1990. 

47. Managit, C. et al., Targeted and sustained drug delivery using PEGylated galactosylated liposomes, 

Int. J. Pharm., 266, 77, 2003. 

48. Maruyama, K., Kennel, S. J., and Huang, L., Lipid composition is important for highly efficient 

target binding and retention of immunoliposomes, Proc. Natl Acad. Sci. U.S.A, 87, 5744, 1990. 

49. Ashwell, G. and Harford, J., Carbohydrate-specific receptors of the liver, Annu. Rev. Biochem., 51, 

531, 1982. 

50. Kuiper, J., The Liver: Biology and Pthobiology, 3rd ed., Raven Press, New York, 1994. 

51. Weitman, S. D. et al., Distribution of the folate receptor GP38 in normal and malignant cell lines and 

tissues, Cancer Res., 52, 3396, 1992. 

52. Yuan, F. et al., Vascular permeability in a human tumor xenograft: Molecular size dependence and 

cutoff size, Cancer Res., 55, 3752, 1995. 

53. Gabizon, A. and Papahadjopoulos, D., Liposome formulations with prolonged circulation time in 

blood and enhanced uptake by tumors, Proc. Natl Acad. Sci. U.S.A, 85, 6949, 1988. 

54. Takino, T. et al., Long circulating emulsion carrier systems for highly lipophilic drugs, Biol. Pharm. 

Bull., 17, 121, 1994. 

55. Webb, M. S. et al., Sphingomyelin-cholesterol liposomes significantly enhance the pharmacokinetic 

and therapeutic properties of vincristine in murine and human tumour models, Br. J. Cancer, 72, 896, 

1995. 

56. Krishna, R. et al., Liposomal and nonliposomal drug pharmacokinetics after administration of liposome-encapsulated 

vincristine and their contribution to drug tissue distribution properties, 

J. Pharmacol. Exp. Theor., 298, 1206, 2001. 

57. Blume, G. and Cevc, G., Liposomes for the sustained drug release in vivo, Biochim. Biophys. Acta, 

1029, 91, 1990. 

58. Gabizon, A. A., Barenholz, Y., and Bialer, M., Prolongation of the circulation time of doxorubicin 

encapsulated in liposomes containing a polyethylene glycol-derivatized phospholipid: Pharmacokinetic 

studies in rodents and dogs, Pharm. Res., 10, 703, 1993. 

59. Gabizon, A. A., Pegylated liposomal doxorubicin: Metamorphosis of an old drug into a new form of 

chemotherapy, Cancer Invest., 19, 424, 2001. 

60. Kawakami, S. et al., Asialoglycoprotein receptor-mediated gene transfer using novel galactosylated 

cationic liposomes, Biochem. Biophys. Res. Commun., 252, 78, 1998. 

61. Kawakami, S. et al., Novel galactosylated liposomes for hepatocyte-selective targeting of lipophilic 

drugs, J. Pharm. Sci., 90, 105, 2001. 

62. Kawakami, S. et al., Biodistribution characteristics of mannosylated, fucosylated, and galactosylated 

liposomes in mice, Biochim. Biophys. Acta, 1524, 258, 2000. 

63. Kawakami, S. et al., Targeted delivery of prostaglandin E1 to hepatocytes using galactosylated 

liposomes, J. Drug Target., 8, 137, 2000. 

64. Hattori, Y. et al., Controlled biodistribution of galactosylated liposomes and incorporated probucol 

in hepatocyte-selective drug targeting, J. Control Release, 69, 369, 2000. 

65. Wang, S. N. et al., Synthesis of a novel galactosylated lipid and its application to the hepatocyteselective 

targeting of liposomal doxorubicin, Eur. J. Pharm. Biopharm., 62, 32, 2006. 

66. Kawakami, S. et al., Mannose receptor-mediated gene transfer into macrophages using novel mannosylated 

cationic liposomes, Gene Theor., 7, 292, 2000. 

67. Kawakami, S. et al., The effect of lipid composition on receptor-mediated in vivo gene transfection 

using mannosylated cationic liposomes in mice, S.T.P. Pharm. Sci., 11, 117, 2001. 

68. Kawakami, S. et al., Effect of cationic charge on receptor-mediated transfection using mannosylated 

cationic liposome/plasmid DNA complexes following the intravenous administration in mice, Pharmazie, 

59, 405, 2004. 



69. Taniyama, T. and Holden, H. T., Direct augmentation of cytolytic activity of tumor-derived macrophages 

and macrophage cell lines by muramyl dipeptide, Cell. Immunol., 48, 369, 1979. 

70. Fox, A. and Fox, K., Rapid elimination of a synthetic adjuvant peptide from the circulation after 

systemic administration and absence of detectable natural muramyl peptides in normal serum at 

current analytical limits, Infect. Immunol., 59, 1202, 1991. 

71. Opanasopit, P. et al., Inhibition of liver metastasis by targeting of immunomodulators using mannosylated 

liposome carriers, J. Control Release, 80, 283, 2002. 

72. Brenner, B. M., Hostetter, T. H., and Humes, H. D., Glomerular permselectivity: Barrier function 

based on discrimination of molecular size and charge, Am. J. Physiol., 234, F455, 1978. 

73. Maack, T. et al., Renal filtration, transport, and metabolism of low-molecular-weight proteins: A 

review, Kidney Int., 16, 251, 1979. 

74. Mihara, K. et al., Disposition characteristics of protein drugs in the perfused rat kidney, Pharm. Res., 

10, 823, 1993. 

75. Mihara, K. et al., Disposition characteristics of model macromolecules in the perfused rat kidney, 

Biol. Pharm. Bull., 16, 158, 1993. 

76. Sawai, K. et al., Renal disposition characteristics of oligonucleotides modified at terminal linkages in 

the perfused rat kidney, Antisense Res. Dev., 5, 279, 1995. 

77. Sawai, K. et al., Disposition of oligonucleotides in isolated perfused rat kidney: Involvement of 

scavenger receptors in their renal uptake, J. Pharmacol. Exp. Theor., 279, 284, 1996. 

78. Kawabata, K., Takakura, Y., and Hashida, M., The fate of plasmid DNA after intravenous injection 

in mice: Involvement of scavenger receptors in its hepatic uptake, Pharm. Res., 12, 825, 1995. 

79. Hashida, M. et al., Pharmacokinetics and targeted delivery of proteins and genes, J. Control Release, 

41, 91, 1996. 

80. Fujita, T. et al., Targeted delivery of human recombinant superoxide dismutase by chemical modification 

with mono- and polysaccharide derivatives, J. Pharmacol. Exp. Theor., 263, 971, 1992. 

81. Mahato, R. I. et al., In vivo disposition characteristics of plasmid DNA complexed with cationic 

liposomes, J. Drug Target., 3, 149, 1995. 

82. Ma, S. F. et al., Cationic charge-dependent hepatic delivery of amidated serum albumin, J. Control 

Release, 102, 583, 2005. 

83. Yamasaki, Y. et al., Pharmacokinetic analysis of in vivo disposition of succinylated proteins targeted 

to liver nonparenchymal cells via scavenger receptors: Importance of molecular size and negative 

charge density for in vivo recognition by receptors, J. Pharmacol. Exp. Theor., 301, 467, 2002. 

84. Fujita, T. et al., Control of in vivo fate of albumin derivatives utilizing combined chemical modification, 

J. Drug Target., 2, 157, 1994. 

85. Yabe, Y. et al., Targeted delivery and improved therapeutic potential of catalase by chemical 

modification: Combination with superoxide dismutase derivatives, J. Pharmacol. Exp. Theor., 

289, 1176, 1999. 

86. Bielinska, A. et al., Regulation of in vitro gene expression using antisense oligonucleotides or 

antisense expression plasmids transfected using starburst PAMAM dendrimers, Nucleic Acids 

Res., 24, 2176, 1996. 

87. Tang, M. X., Redemann, C. T., and Szoka, F. C., In vitro gene delivery by degraded polyamidoamine 

dendrimers, Bioconjug. Chem., 7, 703, 1996. 

88. Sakurai, F. et al., Interaction between DNA–cationic liposome complexes and erythrocytes is an 

important factor in systemic gene transfer via the intravenous route in mice: The role of the neutral 

helper lipid, Gene Theor., 8, 677, 2001. 

89. Mahato, R. I. et al., Physicochemical and pharmacokinetic characteristics of plasmid DNA/cationic 

liposome complexes, J. Pharm. Sci., 84, 1267, 1995. 

90. Zhu, N. et al., Systemic gene expression after intravenous DNA delivery into adult mice, Science, 

261, 209, 1993. 

91. Song, Y. K. et al., Characterization of cationic liposome-mediated gene transfer in vivo by intravenous 

administration, Hum. Gene Theor., 8, 1585, 1997. 

92. Uyechi, L. S. et al., Mechanism of lipoplex gene delivery in mouse lung: Binding and internalization 

of fluorescent lipid and DNA components, Gene Theor., 8, 828, 2001. 

93. Sakurai, F. et al., Therapeutic effect of intravenous delivery of lipoplexes containing the interferonbeta 

gene and poly I: poly C in a murine lung metastasis model, Cancer Gene Theor., 10, 661, 2003. 


58 


94. Yamada, M. et al., Tissue and intrahepatic distribution and subcellular localization of a mannosylated 

lipoplex after intravenous administration in mice, J. Control Release, 98, 157, 2004. 

95. Hattori, Y. et al., Enhancement of immune responses by DNA vaccination through targeted gene 

delivery using mannosylated cationic liposome formulations following intravenous administration in 

mice, Biochem. Biophys. Res. Commun., 317, 992, 2004. 

96. Hara, T. et al., Receptor-mediated transfer of pSV2CAT DNA to a human hepatoblastoma cell line 

HepG2 using asialofetuin-labeled cationic liposomes, Gene, 159, 167, 1995. 

97. Kawakami, S. et al., In vivo gene delivery to the liver using novel galactosylated cationic liposomes, 

Pharm. Res., 17, 306, 2000. 

98. Fumoto, S. et al., Enhanced hepatocyte-selective in vivo gene expression by stabilized galactosylated 

liposome/plasmid DNA complex using sodium chloride for complex formation, Mol. Theor., 

10, 719, 2004. 

99. Hofland, H. E. et al., Folate-targeted gene transfer in vivo, Mol. Theor., 5, 739, 2002. 

100. Reddy, J. A. et al., Folate-targeted, cationic liposome-mediated gene transfer into disseminated 

peritoneal tumors, Gene Theor., 9, 1542, 2002. 


5 

CONTENTS 

Multifunctional Nanoparticles 

for Cancer Therapy 

Todd J. Harris, Geoffrey von Maltzahn, 

and Sangeeta N. Bhatia 

5.1 Introduction ........................................................................................................................... 59 

5.2 Modular Functionalities at the Biosynthetic Interface ......................................................... 60 

5.2.1 Targeting.................................................................................................................... 60 

5.2.2 Imaging Agents ......................................................................................................... 61 

5.2.3 Sensing ...................................................................................................................... 64 

5.2.4 Therapeutic Payloads ................................................................................................ 66 

5.2.5 Remote Actuation...................................................................................................... 69 

5.3 Challenges in Integrating Multiple Functionalities and Future Directions ......................... 70 

References ...................................................................................................................................... 70 


The use of nanoparticles in cancer therapy is attractive for several reasons: they exhibit unique 

pharmacokinetics, including minimal renal filtration; they have high surface-to-volume ratios 

enabling modification with various surface functional groups that home, internalize, or stabilize; 

and they may be constructed from a wide range of materials used to encapsulate or solubilize 

therapeutic agents for drug delivery or to provide unique optical, magnetic, and electrical properties 

for imaging and remote actuation. The topology of a nanoparticle—core, coating, and surface 

functional groups—makes it particularly amenable to modular design, whereby features and 

functional moieties may be interchanged or combined. Although many functionalities of nanoparticles 

have been demonstrated, including some clinically approved drug formulations and imaging 

agents, 3,8 the consolidation of these into multifunctional nanoparticles capable of targeting, 

imaging, and delivering therapeutics is an exciting area of research that holds great promise for 

cancer therapy in the future. 

Figure 5.1 1 schematically depicts a hypothetical multifunctional particle that has been engineered 

to include many features such as the ability to target tumors, evade uptake by the 

reticuloendothelial system (RES), protect therapeutics that can be released on demand, act as 

sensors of tumor responsiveness, and provide image contrast to visualize sites of disease and 

monitor disease progression. Some of these features, such as targeting, leverage biological 

machinery. Others are derived synthetically and enable external probing or manipulation that is 

otherwise not feasible in biological systems. In this chapter, we review both bio-inspired and 

synthetic nanoparticle functionalities that have been used in cancer therapy and address both 

current efforts and future opportunities to combine these into multifunctional devices. 


59

60 

Targeting species 

Therapeutic 

5.2 MODULAR FUNCTIONALITIES AT THE BIOSYNTHETIC INTERFACE 

5.2.1 TARGETING 

Diagnostic 

Actuate/Trigger 

Sensor 


FIGURE 5.1 Schematic depiction of a multifunctional nanoparticle. A hypothetical nanoparticle targets the 

tumor, senses and reports molecular signatures, and delivers a therapeutic in response to an external or 

biological trigger. (From Ruoslahti, E., Cancer Cell, 2, 97–98, 2002.) 

The ability to physically target diseased cells to receive therapeutics while avoiding residual uptake 

in other tissues has long been a goal in cancer therapy. 9–12 The homing of stem cells to a tissue 

niche, or the susceptibility of one cell over another to viral infection, demonstrates that biomolecular 

recognition can be used to direct species to specific extracellular and intracellular sites. The 

microenvironment of the tumor including cell-surface markers, extracellular matrix, soluble 

factors, and proteases, as well as the tumor’s unique architecture and transport properties, may 

be exploited for targeting. 13–17 

Both passive and active targeting have been utilized for nanoparticle delivery. Passive targeting 

relies upon the unique pharmacokinetics of nanoparticles including minimal renal clearance and 

enhanced permeability and retention (EPR) through the porous angiogenic vessels in the tumor. 18,19 

Surface attachment of polymers such as poly(ethylene glycol) (PEG) and poly(ethylene oxide) 

(PEO) enables nanoparticles to avoid uptake by mononuclear phagocytes in the liver, spleen, and 

lymph nodes, thereby improving accumulation in the tumor. 20–22 Active targeting relies on liganddirected 

binding of nanoparticles to receptors expressed in the tumor. Binding of ligands to the 

vasculature can occur immediately, as it is directly accessible to nanoparticles circulating in the 

blood. Over longer time periods, particles extravasate into the tissues where receptors expressed on 

cancer cells and in the interstitium may be used for localization. 23–25 

Many candidate tumor markers have been described, some of which bind known ligands such 

as arginine–glycine–aspartic acid (RGD)-binding Vb3 and Vb5 integrins expressed on the surface of 

angiogenic blood vessels, and folic acid-binding receptors on the surface of cancer cells. These and 

others have been attached to the surface of various nanoparticle cores to deliver them to 

tumors. 17,26–28 Monoclonal antibodies have also been used extensively for targeting. These can 

be isolated with high affinity for tumor markers and are useful for targeting receptors of unknown or 

low affinity ligands. 29,30 Novel screens for discovering tumor homing ligands have been developed 

using phage and bacterial display as well as libraries of aptamers, peptides, polymers, and small 

molecules. 31,32 These techniques may be used to isolate targeting ligands, even when their target 


Multifunctional Nanoparticles for Cancer Therapy 61 

receptor is unknown. For example, the 34 amino acid, cationic peptide F3, which has been used to 

deliver quantum dots to tumor endothelium, was uncovered initially by a blind-page display screen 

in a breast cancer xenograft model and later found to bind cell surface nucleolin expressed on tumor 

endothelium and cancer cells. 6,33,34 

Although extracellular targeting to the tumor is sufficient for many modes of imaging and drug 

delivery, intracellular delivery of nanoparticles into the cytosol is essential for some applications. 

For example, nanoparticles carrying membrane-impermeable cargo that perform their function in 

the cytosol, such as siRNA, antisense DNA, peptides, and other drugs, are minimally effective if 

delivered extracellularly or sequestered in the endosome. 35 Protein and peptide motifs capable of 

translocating nanoparticles into the cytoplasm have been borrowed from mechanisms of viral 

transfection. Two important classes of translocating domains include polycationic sequences and 

membrane fusion domains. Attaching the short polycationic sequence of HIV’s TAT protein, amino 

acid residues 48–57, to a nanoparticle facilitates its adsorption on a cell surface and subsequent 

internalization into the cell. 36,37 This peptide has been used to internalize dextran coated iron-oxide 

nanoparticles into T-cells in vitro, which were subsequently used to monitor T-cell trafficking in 

tumors with MRI. 38 Use of this peptide for intracellular delivery in vivo is limited by the adverse 

effect that polycationic sequences have on nanoparticle circulation time and RES uptake. 39 The 

amphiphilic domain derived from the N-terminus of the influenza protein hemagluttinin (HA2) is a 

membrane fusion peptide that destabilizes the endosome at low pH and facilitates viral escape into 

the cytosol. 40 Variations of this peptide with improved infectivity have also been synthesized. 41 

Influenza-derived peptides have been used to enhance the delivery of liposomes as well as 100 nm 

poly-L-lysine particles. Although the peptide modification of these particles improves endosomal 

escape over unmodified particles, the transfection efficiency still remains well below that of intact 

viruses. 42,43 

Another level of targeting can occur after translocation of nanoparticles into the cytosol to 

direct nanoparticles to specific sub-cellular structures. Using peptide localization sequences, fluorescent 

quantum dots have been targeted to the nucleus and the mitochondria (Figure 5.2). 2 Several 

other localization sequences exist and could be used to traffic nanoparticles to the endoplasmic 

reticulum, golgi apparatus, or peroxisomes. Although work in this area has been focused on 

organelle labeling, the potential for delivering therapeutic nanoparticles to sub-cellular structures 

is possible. Such nanoparticles could sense sub-cellular aspects of disease or specifically intervene 

for more potent treatment or eradication of cancer cells (i.e., free-radical-mediated mitochondrial 

damage to induce apoptosis). 

5.2.2 IMAGING AGENTS 

Imaging cancer is crucial for guiding decisions about treatment and for monitoring the efficacy of 

administered therapies. The use of nanoparticles for image contrast and enhancement has enabled 

improvements in cancer imaging by conventional modalities, such as magnetic resonance imaging 

(MRI) and ultrasound, and has also established new techniques such as optical-based imaging for 

cancer detection. 39,44,45 Targeted imaging agents that can identify specific biomarkers have the 

potential to improve detection, classification, and treatment of cancer with minimal invasiveness 

and reduced costs. 

The use of nanoparticles in cancer imaging has already demonstrated clinical efficacy in 

detecting liver cancer and staging lymph node metastasis noninvasively. 3,46 Superparamagnetic 

iron-oxide nanoparticles disrupt local magnetic field gradients in tissues, causing a detectable signal 

void in MRI. Dextran coated iron-oxide nanoparticles administered intravenously get phagocytosed 

by normal macrophages of the liver and lymph and the failure of these tissues to darken after ironoxide 

administration identifies invading cancer cells. Directly targeting these magnetic nanoparticles 

to cancer cells has also been demonstrated. For example, herceptin mAb and folic acid on the 

surface of iron-oxide nanoparticles enable MRI-based molecular imaging of their respective targets 


62 

QD+NLS 

70kD Dextran Merge 

QD+MLS Mitotracker 

Merge 


FIGURE 5.2 (See color insert following page 522.) Labeling of intracellular targets with peptide labeled 

quantum dots. Quantum dots (QD) modified with PEG and a nuclear localization sequence (NLS) or a 

mitochondrial localization sequence (MLS) were shown to target the nucleus or mitochondria of cells respectively. 

Seventy kilo Dalton PEG distributed in the cytoplasm contrasts nuclear localization while MitoTracker 

colocalizes with mitochondrial localization. (From Derfus, A. M., Chan, W. C. W., and Bhatia, S. N., 

Advanced Materials, 16, 961, 2004.) 

in tumors. 47,48 Other nanoparticle cores including dendrimers, micelles, and liposomes modified 

with paramagnetic gadolinium have also been used for tumor targeted MRI contrast. 48–50 

Gold nanoshells offer a promising alternative to MRI probes by providing contrast for optical 

imaging. 45 These nanoparticles are constructed from a dielectric core (silicon) and a metallic 

conducting shell (gold). By varying the dimension of the core and shell, the plasmon resonance of 

these particles can be engineered to either absorb or scatter wavelengths of light, from UV to infrared. 

Particles that are tailored to scatter light in the near-infrared, where tissues have minimal absorbance, 

have been used to enhance imaging modalities such as reflectance confocal microscopy and optical 

coherence tomography (OCT). 51 Although the penetration of optical techniques does not approach 

that of CT or MRI, imaging features is possible at depths of a few centimeters. Gold colloids have also 

been used for optical contrast, but these lack the inherent tunability of nanoshells. The conjugation of 

optical contrast agents to antibodies has been used for the molecular imaging of the EGFR receptor on 

early cervical precancers and for Her2C breast carcinoma cells in mice. 52,53 

Fluorescent nanoparticles offer another useful tool to enhance optical detection. These probes 

are identified easily in microscopy and are useful for tracking the biodistribution of nanoparticles 

in experimental models. Fluorescent semiconductor nanocrystals, quantum dots, have been used 

to show ligand-mediated nanoparticle targeting to distinct features in the tumor. 6 Three different 

phage display-derived peptides were used to specifically target these nanocrystals to tumor blood 

vessels, tumor lymphatics, or lung endothelium (Figure 5.3). Functionalization of these nanoparticles 

with PEG eliminated detectable accumulation in RES organs, including the liver 

(Figure 5.4). Quantum dots have a distinct advantage over conventional fluorophores because 



(a) 

Prostate 

FIGURE 5.3 (See color insert following page 522.) Noninvasive detection of lymph node metastasis with 

iron-oxide nanoparticles and MRI. (a) A three-dimensional reconstruction using nanoparticle-enhanced MRI 

of the prostate. Metastatic lymph nodes are in red and normal nodes are in green. (b) Conventional MRI image 

shows similar signal intensity from two adjacent nodes. (c) Nanoparticle-enhanced MRI shows a decreased 

signal in the normal node from macrophage uptake (thick arrow), but not in the metastatic node (thin arrow). 

(From Harisinghani, M. G. and Weissleder, R., Plos Medicine, 1, 202–209, 2004.) 

of their size-tunable excitation and emission profiles, narrow bandwidths, and high photo-stability. 

54,55 Using nanocrystals that fluoresce in the near-infrared could extend their utility to clinical 

settings, 56 though a key limitation has been their potential toxicity because they are formulated 

from heavy metals. 57 Efforts to make these of nontoxic materials are ongoing (Figure 5.4). 

Alternative fluorescent nanoparticle probes have been developed including fluorescently tagged 

dendrimers and fluorophore-embedded silica nanoparticles. 58–60 

Nanoparticle formulations that provide contrast for other imaging modalities including ultrasound 

and CT have been described. Perfluorocarbon emulsion nanoparticles composed of lipidencapsulated 

perfluorocarbon liquid, about 250 nm diameter, are effective in giving echo contrast. 61 

Air-entrapping liposomes formulated from freeze-drying techniques have also been developed to 

give ultrasound contrast. 62,63 These agents passively distribute in RES organs and areas of angiogenenis 

enabling enhanced imaging of these features. Bismuth sulfide nanoparticles can be used as 

contrast agents for CT imaging, giving blood pool contrast similar to that of iodine but at lower 

concentrations. These can also be used to image lymph nodes after phagocytic uptake. 64 As with 

other imaging modalities, the attachment of appropriate ligands to nanoparticle-based CT and 

ultrasound contrast agents could be used for molecular imaging. 

An interesting extension of the imaging agents described above is their combination into 

multimodal imaging nanoparticles. The combination of fluorescence and magnetic properties 

within a single particle has been used for dual optical and MRI imaging. Fluorescence is used to 

determine the localization of targeted imaging agents down to specific micro-structures inside and 

outside cells and magnetic domains provide three-dimensional whole-body imaging capabilities 

with MRI. 59,65 These magneto-fluorescent nanoparticles could be effective for image guided 

surgeries where MRI is used to locate cancer in the body and fluorescence is used to more precisely 


(b) 

(c)

64 

(a) 

(b) 

Peptide 

Peptide 

Peptide 

delineate tumor borders during resection. Other dual-imaging probes have been described 

including: perfluorocarbon emulsions tagged with gadolinium for combined ultrasound and MRI. 66 

5.2.3 SENSING 

Peptide 

ZnS 

CdSe 

Peptide 

Peptide 

Peptide 

Peptide 

Peptide 

Peptide 

PEG 

ZnS 

PEG 

Functionalities that undergo chemical alterations in response to enzymatic activity or other 

properties such as pH or oxygen could be used as sensors to report information about the status of 

the tumor or efficacy of treatment. Many nanoparticle-based sensors that respond to biological 

triggers including proteases, DNAses, proteins, peroxidase, pH, and others have been demonstrated 

in vitro. 67–72 These generally rely on assembly or disassembly of inorganic nanocrystals 

including: gold nanoparticles, nanoshells, or nanorods, which undergo a shift in their plasmon 

resonance when aggregated; iron-oxide nanoparticles have enhanced T2 relaxivity when clustered; 

and fluorescent quencher-based nanoparticle systems that dequench after triggered release. 

A system using cleavable polymeric shielding of self-assembling nanoparticles has been 

proposed as a mechanism for translating nanoparticle-based enzyme sensors to in vivo use 

(Figure 5.5). 73 Self-assembling, complementary iron-oxide nanoparticles are rendered latent with 

PEG polymers linked to the nanoparticle surface by protease-cleavable substrates that serve to both 

inhibit assembly and stabilize the particles in serum. Upon proteolytic removal of PEG polymers by 

MMP-2 expressing cancer cells, nanoparticles assemble and acquire amplified magnetic properties 

that can be detected with MRI. In the future, similar to thrombin-driven self-assembly of fibrin and 

platelets at sites of endothelial injury, this system may allow the hyper-active proteolytic processes of 

cancers to drive the self-assembly of nanoparticles in regions of cancer angiogenesis, invasion, and 

metastasis in vivo. Due to its modular design, this system can easily be modified for a number of 

detection schemes by substituting the complementary binding pairs, cleavable substrates 

Peptide 

PEG CdSe Peptide 

FIGURE 5.4 (See color insert following page 522.) Targeting quantum dots (QDs) to site-specific endothelium 

with phage display-derived peptides. (a) Schematic representation of co-injected red and green 

quantum dots that home to tumor and lung vasculature respectively after intravenous injection. (b) Schematic 

representation of peptide-coated QDs and peptide-coated PEG-QDs. (c) QDs labeled with the tumor endothelium 

homing peptide F3 co-localize with a blood vessel marker. (d) QDs labeled with the tumor lymphatic 

homing peptide Lyp-1 highlight the endothelium but do not colocalize with a blood vessel marker. (From 

Akerman, M. E., Chan, W. C. W., Laakkonen, P., Bhatia, S. N., and Ruoslahti, E., Proceedings of the National 

Academy of Sciences of the United States of America, 99, 12617–12621, 2002.) 


PEG 

(c) 

(d) 

Nanotechnology for Cancer Therapy

(a) 

(b) 

0.03 

0.02 

Δe 

0.01 

0.00 


0 

No MMP-2 

Scmbl cntrl 

MMP-2 

30 60 

Diffusivity 

'Latent' nanoparticles Proteolytically actuated self-assembly 

90 

t /min 

120 150 180 

(c) (d) 

10 kDa PEG 

MMP-2 

MMP-2 

Nanoassemblies exhibit: 

Magnetic susceptibility 

T2 relaxativity 

No MMP-2 

32 

particle 10 

concn./pM 

3.2 

0 

85 170 340 680 1360 

MMP-2/ng/ml 

400 

300 

200 

T2/ms 

FIGURE 5.5 (See color insert following page 522.) Protease-triggered self-assembling nanoparticles with polymer-shielded coatings. (a) Schematic representation of 

nanoparticles that self-assemble after protease-mediated cleavage of PEG chains reveals complementary moieties (neutravidin and biotin). (b) Iron-oxide nanoparticles 

with cleavable linkers assemble in the presence of MMP-2, as measured by changes their light extinction, while particles with noncleavable scrambled peptides do not. (c) 

Atomic force micrographs of particles incubated with MMP-2 shows detectable aggregation (scale bars are 500 nm). (d) T2 maps of particles in solution using a 4.7 T 

MRI demonstrate enhanced T2 relaxivity for increasing concentrations of MMP-2. 

100 

Multifunctional Nanoparticles for Cancer Therapy 65

66 

(e.g., glycans, lipids, oligonucleotides), or multivalent nanoparticle cores (e.g., gold, quantum dot, 

dendrimer). 

5.2.4 THERAPEUTIC PAYLOADS 


The use of nanoparticulate drug carriers can address many critical challenges in drug delivery 

including: improving drug solubility and stability; extending drug half-lives in the blood; reducing 

adverse effects in nontarget organs; and concentrating drugs at the disease site. 74 Drugs may be 

dispersed in a matrix, encapsulated in a vesicle, dissolved in a hydrophobic core, or attached to the 

surface of a nanoparticle. Several nanoparticle-based drug delivery systems including liposomes, 

polymeric nanoparticles, dendrimers, ceramic based carriers, micelles, and others have been used to 

carry small molecule, peptide, and oligonucleotide therapeutic agents. 24,75 Many promising 

anti-cancer drugs fail to make it to the clinic because of poor solubility or high collateral toxicity 

at therapeutic levels, thus motivating the need for these carriers in cancer therapy. 

Liposomes have been the most extensively utilized nanoparticle-based carriers for delivering 

anti-cancer drugs. First described decades ago, these submicron-sized carriers consist of amphiphilic 

lipids assembled to form vesicles that can encapsulate drugs. 76 Liposome-encapsulated doxorubicin 

is a clinically approved nanoparticle formulation used for chemotherapy. 77 The surface of this 

nanocarrier is PEGylated to reduce rapid uptake by phagocytic cells and extend the drug circulation 

time for better therapeutic efficacy. Several other liposome-encapsulated chemotherapeutic drugs 

have been described, with many in clinical trials. 78 Active targeting of these liposomes through the 

attachment of antibodies and various ligands has also been demonstrated. 30,79 Drug loaded liposomes 

with encapsulated or surface-functionalized gadolinium or fluorophores have been used to simultaneously 

image tumors during nanoparticle-targeted drug delivery. 80–82 

Biodegradable polymer nanocarriers have also been investigated as a means of encapsulating 

drugs and releasing them over time. Both poly dl-lactide co-glycolide (PLGA) and polylactide (PLA) 

nanoparticles have been formed that immobilize drugs dispersed in their matrix and release them 

upon degradation. 24,83 Other polymers, including polyethyleneimine (PEI), polylysine, and cyclodextrin-containing 

polymers, are used to condense DNA or RNA into nanoparticle carriers that can 

be targeted to cancer cells for gene or siRNA delivery. 35 Polymeric micelles consist of amphiphilic 

block copolymers that self-assemble into a water-soluble nanoparticle with a hydrophobic core. 

These can be used to encapsulate water-insoluble drugs such as doxorubicin and adriamycin and 

targeted to tumors. 84–86 Polymersomes are another variation of polymer-based nanoparticulate 

vesicles that self-assemble from amphiphilic block copolymers. 87 These have been used to encapsulate 

doxorubicin with well-controlled release over several days. 88,89 

Another class of nanoparticle-based drug carriers are dendrimers. These consist of a network of 

branching chemical bonds around an inner core. One of the more popular dendrimers, polyamidoamine 

dendrimers (PAMAMs), are nonimmunogenic, water-soluble, and possess terminal amine 

functional groups for conjugation of a variety of surface moieties. 90 Their inner core can been used 

to encapsulate anti-cancer drugs such as doxorubicin and methotrexate. 91 Drugsmayalsobe 

conjugated to the dendrimer surface along with ligands for targeting. 92,93 A dendrimer functionalized 

with FITC, folic acid, and methoxetrate has been synthesized to have imaging, targeting, and 

drug delivery capabilities (Figure 5.6). 4,7,94 The synthesis of these conjugates in a scalable and 

reproducible manner has been described for potential clinical applications. 4 

Other nanoparticulate carriers, including nanoemulsions, drug nanocrystals, and polyelectrolyte 

carriers, have been developed. Nanoemulsions are formed by dissolving a drug in a lipid, 

cooling the solution under high pressure, and using homogenization to form solid nanoparticle lipid 

carriers at body temperature. Homogenization techniques can also be used to form crystalline 

nanosuspensions of drugs. 74 These formulations increase drug solubility and control release 

kinetics of the drug in the blood and at the tumor site. Polyelectrolyte carriers formed by the 


H 2 N 

N 

N 

NH 2 

H 

CH 3 

C 

HN 

O 

5 

HO CH O 

C 

O 

MTX 

O HO 

HO 

HO 

HO HOHO 

HO 

OH 

N 

N 

15 

O 

O 

C 

OH 

OH 

H 

N C 

N 

H 

O 

NH 

G5 

N 

H 

O 

C 

CH3 AC 

82 

HN 

C S 

HO 

HN 

O 

COOH 

4 

FITC 

(a) (b) 

FA 

O 

H 

H 

OH 

NH 2 

4 

G5-FI 

G5-FI-PA 

FIGURE 5.6 Multifunctional dendrimers for targeting, imaging, and drug delivery. (a) Schematic of a multifunctional dendrimer labeled with fluorescein (FITC) for 

imaging, folic acid (FA) for targeting, methotrexate (MTX) for therapeutic delivery, and alcohol (OH) and acetylated (Ac) moieties for particle stabilization. (b) Confocal 

images of FITC-labeled dendrimers incubated over cells with and without a targeting antibody, anti-PSMA (PA). (From Majoros, I. J., Thomas, T. P., Mehta, C. B., and 

Baker, J. R., Journal of Medicinal Chemistry, 48, 5892–5899, 2005; Thomas, T. P. et al., Biomacromolecules, 5, 2269–2274, 2004.) 


1 

2 

Multifunctional Nanoparticles for Cancer Therapy 67

68 

layer-by-layer absorption of polycationic and polyanionic moieties can be used to encapsulate 

therapeutic cargo, particularly larger agents such as peptides and oligonucleotides. 95 

A clever combination of drug-release modalities was recently demonstrated by the creation of 

a dual drug-release nanoparticle having a PLGA polymer core encapsulating doxorubicin and a 

PEG-lipid block copolymer shell loaded with the combretastatin (Figure 5.7). 5 The lipophilic 

anti-angiogenesis drug, combretastatin, intercalates in the nanoparticle membrane and releases 

rapidly upon association with tumor endothelial cells, while the slower-releasing doxorubicin 

increases cytotoxic killing of tumor cells for a prolonged time after the vasculature shuts down. 

This novel system demonstrates the feasibility of integrating multiple functionalities of drug 

delivery on a single nanoparticle to enhance therapeutic efficacy. 

Total drug released 

(a) (c) 

(b) (d) 

Tumor volume 

8 

6 

4 

2 

0 

0 20 40 60 80 100 120 

Time (hours) 

7500 

5000 

2500 


Veh 

C 

N 

N+C 

NC 

0 

9 10 11 12 13 

Days 

Combretastatin (×10 2 μg) 

Doxorubicin (μg) 

140 160 

FIGURE 5.7 Dual drug-release nanoparticle for combined anti-angiogenisis and anti-cancer treatment. (a) 

Scanning electron micrograph showing nanocores prepared from doxorubicin-coupled PLGA. The nanocores 

are encapsulated inside a lipid coat, which is also loaded with an anti-angiogenesis agent. (b) Cross section of a 

nanocell with the dark nanocore. The lipid coat is surface-modified through pegylation, which confers stealth 

characteristics to the nanocell from the RES. (c) The composition of the nanocell enables a spatiotemporal 

release of the two agents in an acidic pH mimicking the tumor environment, as shown in the graph. (d) In vivo 

studies using F10 melanoma clearly show that the spatiotemporal release from the nanocells (NC) achieves 

better outcome than the doxorubicin-loaded nanocore or the lipid-entrapped combretastatin (C) alone, or 

combinations of both (N+C). 5 


14


5.2.5 REMOTE ACTUATION 

Temporal and spatial control of therapeutic administration is important for eliminating off-target 

toxicity and achieving optimal delivery. Temporally controlled release profiles may be designed into 

nanoparticle carriers mentioned previously and spatial control can be improved with targeting. 

However, off-target effects, including eventual accumulation of nanoparticles in RES organs, limit 

many aspects of these methods of control. The ability to trigger the therapeutic activity of administered 

nanoparticles remotely could be a valuable tool for localizing treatments to a diseased site. 

Many inorganic nanocrystals and nanoemulsions used for imaging contrast absorb electromagnetic 

or ultrasonic energy that can also be used to remotely heat or trigger drug delivery. 

Thermal ablation of tumors by nanoparticles that absorb external energy has been demonstrated 

both with iron-oxide nanoparticles and gold nanoshells. Superparamagnetic iron-oxide nanoparticles 

under the influence of an alternating electromagnetic (EM) field heat by Brownian relaxation, 

where heat is generated by the rotation of particles in the field, and Neel relaxation, where the 

magnetic domains are moved away from their easy axis with the resultant energy being deposited as 

heat in the solution. 96,97 Nanoparticle concentrations of 0.1–1% are required to achieve critical 

temperatures for tumor ablation. 98,99 Ongoing work to increase the absorption of magnetic nanoparticles 

using clinically safe RF frequencies and to increase the concentration of particles that can 

be targeted to the tumor may extend the utility of this technique. Alternatively, near-infraredabsorbing 

gold nanoshells targeted to the tumor can be used to thermally ablate the cancer 

cells upon illumination with a high intensity laser. 100,101 This technique can be applied to solid 

tumors in close proximity to the skin, but cannot be applied to deeper lesions because of tissue 

absorbance. 53,100,101 By synthesizing nanoshells with a plasmon resonance that has both absorption 

and scattering profiles, these nanoparticles may be capable of both heating and imaging 

tumors. 53 

Remotely-triggered release of a therapy by heating is a promising extension of the use of 

nanoparticles that can absorb external energy. An example of this has been demonstrated with a 

model drug linked to an iron-oxide nanoparticle via a heat-labile tether that is released and diffuses 

into the peripheral tissue after irradiation with RF energy. 98 By modifying the susceptibility of the 

linker, it is possible to tune the release profile over a range of temperatures and to enable repeated 

administrations. The iron core of these drug-releasing nanoparticles can be used simultaneously for 

imaging with MRI. Additionally, the magnetic properties of these nanoparticles can be manipulated 

by magnetic field gradients to target sites near externally- or internally-placed magnets. 102 

Drug activation using EM energy has been explored extensively with photodynamic therapy 

(PDT). PDT agents, when irradiated by light, produce reactive oxygen species that are toxic to cells. 

Agents such as porphyrins have been conjugated to various nanoparticle cores including dendrimers, 

liposomes, and polymers. 103,104 When excited by light, these nanoparticles can produce 

enough reactive oxygen species to kill tumor cells. 60 The inherent fluorescent properties of 

many PDT agents enable simultaneous imaging with therapeutic delivery. A multifunctional nanoparticle 

platform combining MRI contrast and photodynamic therapy has been used to target, 

image, and treat brain cancer in a rat model. 105 In the future, integrating these nanoparticles 

with peptides capable of targeting tumors and subcellularly localizing them to the nuclei or mitochondria 

of tumor cells may enhance the therapeutic efficacy of these treatments. 

Other forms of externally applied energy such as ultrasound and x-ray radiation provide 

alternative mechanisms to achieve remote actuation. Acoustic energy has been shown to 

enhance the delivery of lipid drugs from a perfluorocarbon emulsion targeted to cell membranes 

and from doxorubicin-loaded polymeric micelles. 106,107 Atomically dense nanoparticles have been 

shown to increase the absorption of x-ray radiation, enhancing their destructive effect in 

surrounding tissue. 108 There is potential for simultaneous imaging and therapeutic delivery with 

these particles also. 


70 

5.3 CHALLENGES IN INTEGRATING MULTIPLE FUNCTIONALITIES AND 

FUTURE DIRECTIONS 

Although remotely actuated nanoparticle cores such as iron-oxide and metal nanoshells naturally 

lend themselves to dual-imaging and therapeutic applications, the combination of imaging and 

other functionalities using other nanoparticle cores can be challenging. There are inherent trade-offs 

when combining many functional groups into one nanoparticle. In many cases, a limited number of 

attachment sites are available on the particle surface, making it difficult to couple several functional 

groups in sufficient concentration for each to function. Moreover, some groups may interact to 

sterically shield or alter the activity of one another when combined in close proximity. Multiple 

functional moieties on a nanoparticle may also reduce colloidal stability or adversely affect its 

in vivo pharmacokinetics. With significant characterization and fine tuning, dendrimers that 

combine targeting, imaging, and therapeutic moieties on their surface have been synthesized 

successfully. 4 Similar efforts will be necessary to achieve other multifunctional nanoparticles 

with decorated-surface moieties. 

An alternative strategy to consolidate multiple functionalities onto a single particle is to use 

core-shell architecture. In this case, an outer shell with one functionality, such as targeting, may be 

unveiled to reveal an inner core that performs a secondary function such as endosomal escape or 

drug release. This has been demonstrated with the conjugation of targeting moieties or protective 

PEG groups on the surface of dendrimers or polymers via acid-labile chemistries that degrade in the 

lower pH of the endosome and unveil endosomal escape mechanisms on the particle core. 109,110 

This has also been demonstrated with protease-cleavable linkers that release protective polymers on 

the surface of complementary nanoparticles to initiate their self-assembly. 73 

The synthesis of nanoparticles with polar domains is another strategy that could be used to 

incorporate multiple functionalities on a single particle. Janus nanoparticles—named for Janus, the 

Roman God of doorways typically depicted with faces on the front and back of his head—have been 

engineered with two chemically distinct hemispheres or surfaces. These nanoparticles may be 

spherical (with opposing faces of unique composition), dumbbell-shaped (with two equal-sized 

spheres linked together), snowman-shaped, and may have other morphologies as well. 95,111 The 

creation of nanoparticles with spatially separated chemical domains is a step towards replicating the 

controlled polarity exhibited in nature across many length scales. Separate hemispheres may be 

used to isolate and organize functional domains on nanoparticles such that they may simultaneously 

carry targeting molecules, endosomal escape domains, sensing moieties, hydrophilic and hydrophobic 

therapeutics, or contrast agents that otherwise might be mutually inhibitory if randomly 

incorporated. Moreover, there may be specific applications for which the polarity and anisotropy of 

Janus nanoparticles have benefit, such as real-time detection of oriented binding events, targeted 

bridging of multiple components at a tumor cell, directed drug delivery, or guided self-assembly. 

Although there have been many exciting advances in the application of nanoparticles for cancer 

imaging and treatment, the true power of these materials will be in their ability to interact with 

disease processes intelligently. The modular design of functionalities that target, sense, signal, and 

treat and the ongoing efforts to consolidate these into single nanoparticle platforms is one way in 

which such ‘smart’ materials are being developed. The further elucidation of complex biological 

processes in tumorogenesis, the discovery of nanomaterials with other novel properties, and the 

consolidation of biological and synthetic machinery in these materials in new and elegant ways are 

key factors that will determine their future success in cancer therapy. 

REFERENCES 


1. Ruoslahti, E., Antiangiogenics meet nanotechnology, Cancer Cell, 2, 97–98, 2002. 

2. Derfus, A. M., Chan, W. C. W., and Bhatia, S. N., Intracellular delivery of quantum dots for live cell 

labeling and organelle tracking, Advanced Materials, 16, 961–966, 2004. 



3. Harisinghani, M. G. and Weissleder, R., Sensitive, noninvasive detection of lymph node metastases, 

Plos Medicine, 1, 202–209, 2004. 

4. Majoros, I. J., Thomas, T. P., Mehta, C. B., and Baker, J. R., Poly(amidoamine) dendrimer-based 

multifunctional engineered nanodevice for cancer therapy, Journal of Medicinal Chemistry, 48, 

5892–5899, 2005. 

5. Sengupta, S. et al., Temporal targeting of tumour cells and neovasculature with a nanoscale delivery 

system, Nature, 436, 568–572, 2005. 

6. Akerman, M. E., Chan, W. C. W., Laakkonen, P., Bhatia, S. N., and Ruoslahti, E., Nanocrystal 

targeting in vivo, Proceedings of the National Academy of Sciences of the United States of America, 

99, 12617–12621, 2002. 

7. Thomas, T. P. et al., In vitro targeting of synthesized antibody-conjugated dendrimer nanoparticles, 

Biomacromolecules, 5, 2269–2274, 2004. 

8. Gordon, A. N. et al., Recurrent epithelial ovarian carcinoma: A randomized phase III study of pegylated 

liposomal doxorubicin versus topotecan, Journal of Clinical Oncology, 19, 3312–3322, 2001. 

9. Ruoslahti, E., Drug targeting to specific vascular sites, Drug Discovery Today, 7, 1138–1143, 2002. 

10. Allen, T. M., Charrois, G. J. R., and Sapra, P., Recent advances in passively and actively targeted 

liposomal drug delivery systems for the treatment of cancer, Abstracts of Papers of the American 

Chemical Society, 226, U458, 2003. 

11. Allen, T. M. and Cullis, P. R., Drug delivery systems: Entering the mainstream, Science, 303, 

1818–1822, 2004. 

12. Wickham, T. J., Targeting adenovirus, Gene Therapy, 7, 110–114, 2000. 

13. Ruoslahti, E. and Rajotte, D., An address system in the vasculature of normal tissues and tumors, 

Annual Review of Immunology, 18, 813–827, 2000. 

14. Jain, R. K., Delivery of molecular and cellular medicine to solid tumors, Advanced Drug Delivery 

Reviews, 46, 149–168, 2001. 

15. Jain, R. K., Delivery of molecular and cellular medicine to solid tumors, Journal of Controlled 

Release, 53, 49–67, 1998. 

16. Jain, R. K., Delivery of molecular and cellular medicine to solid tumors, Microcirculation—London, 

4, 3–23, 1997. 

17. Satchi-Fainaro, R. et al., Targeting angiogenesis with a conjugate of HPMA copolymer and 

TNP-470, Nature Medicine, 10, 255–261, 2004. 

18. Matsumura, Y. and Maeda, H., A new concept for macromolecular therapeutics in cancerchemotherapy—mechanism 

of tumoritropic accumulation of proteins and the antitumor agent 

Smancs, Cancer Research, 46, 6387–6392, 1986. 

19. Tabata, T., Murakami, Y., and Ikada, Y., Tumor accumulation of poly(vinyl alcohol) of different 

sizes after intravenous injection, Journal of Controlled Release, 50, 123–133, 1998. 


Theory to practice, Pharmacological Reviews, 53, 283–318, 2001. 

21. Moghimi, S. M. and Hunter, A. C., Recognition by macrophages and liver cells of opsonized 

phospholipid vesicles and phospholipid headgroups, Pharmaceutical Research, 18, 1–8, 2001. 

22. Nicolazzi, C. et al., Anionic polyethyleneglycol lipids added to cationic lipoplexes increase their 

plasmatic circulation time, Journal of Controlled Release, 88, 429–443, 2003. 

23. Ruoslahti, E., Specialization of tumour vasculature, Nature Reviews Cancer, 2, 83–90, 2002. 

24. Sahoo, S. K. and Labhasetwar, V., Nanotech approaches to delivery and imaging drug, Drug 

Discovery Today, 8, 1112–1120, 2003. 

25. Wickline, S. A. and Lanza, G. M., Nanotechnology for molecular imaging and targeted therapy, 

Circulation, 107, 1092–1095, 2003. 

26. Stella, B. et al., Design of folic acid-conjugated nanoparticles for drug targeting, Journal of Pharmaceutical 

Sciences, 89, 1452–1464, 2000. 

27. Lockman, P. R. et al., Brain uptake of thiamine-coated nanoparticles, Journal of Controlled Release, 

93, 271–282, 2003. 

28. Dubey, P. K., Mishra, V., Jain, S., Mahor, S., and Vyas, S. P., Liposomes modified with cyclic RGD 

peptide for tumor targeting, Journal of Drug Targeting, 12, 257–264, 2004. 

29. Kirpotin, D. et al., Sterically stabilized Anti-HER2 immunoliposomes: Design and targeting to 

human breast cancer cells in vitro, Biochemistry, 36, 66–75, 1997. 


72 


30. Li, L. Y. et al., A novel antiangiogenesis therapy using an integrin antagonist or anti-FLK-1 antibody 

coated Y-90-labeled nanoparticles, International Journal of Radiation Oncology Biology Physics, 

58, 1215–1227, 2004. 

31. Farokhzad, O. C. et al., Nanopartide-aptamer bioconjugates: A new approach for targeting prostate 

cancer cells, Cancer Research, 64, 7668–7672, 2004. 

32. Allen, T. M., Sapra, P., Moase, E., Moreira, J., and Iden, D., Adventures in targeting, Journal of 

Liposome Research, 12, 5–12, 2002. 

33. Christian, S. et al., Nucleolin expressed at the cell surface is a marker of endothelial cells in 

angiogenic blood vessels, Journal of Cell Biology, 163, 871–878, 2003. 

34. Porkka, K., Laakkonen, P., Hoffman, J. A., Bernasconi, M., and Ruoslahti, E., A fragment of the 

HMGN2 protein homes to the nuclei of tumor cells and tumor endothelial cells in vivo, Proceedings 

of the National Academy of Sciences of the United States of America, 99, 7444–7449, 2002. 

35. Pack, D. W., Hoffman, A. S., Pun, S., and Stayton, P. S., Design and development of polymers for 

gene delivery, Nature Reviews Drug Discovery, 4, 581–593, 2005. 

36. Frankel, A. D. and Pabo, C. O., Cellular uptake of the tat protein from human immunodeficiency 

virus, Cell, 55, 1189–1193, 1988. 

37. Wadia, J. S., Stan, R. V., and Dowdy, S. F., Transducible TAT-HA fusogenic peptide enhances escape of 

TAT-fusion proteins after lipid raft macropinocytosis, Nature Medicine, 10, 310–315, 2004. 

38. Kircher, M. F. et al., In vivo high resolution three-dimensional imaging of antigen-specific cytotoxic 

T-lymphocyte trafficking to tumors, Cancer Research, 63, 6838–6846, 2003. 

39. Weissleder, R., Bogdanov, A., Neuwelt, E. A., and Papisov, M., Long-circulating iron-oxides for 

Mr-imaging, Advanced Drug Delivery Reviews, 16, 321–334, 1995. 

40. Wagner, E., Application of membrane-active peptides for nonviral gene delivery, Advanced Drug 

Delivery Reviews, 38, 279–289, 1999. 

41. Plank, C., Oberhauser, B., Mechtler, K., Koch, C., and Wagner, E., The influence of endosomedisruptive 

peptides on gene-transfer using synthetic virus-like gene-transfer systems, Journal of 

Biological Chemistry, 269, 12918–12924, 1994. 

42. Mastrobattista, E., Crommelin, D. J. A., Wilschut, J., and Storm, G., Targeted liposomes for delivery of 

protein-based drugs into the cytoplasm of tumor cells, Journal of Liposome Research, 12, 57–65, 2002. 

43. Kakudo, T. et al., Transferrin-modified liposomes equipped with a pH-sensitive fusogenic peptide: 

An artificial viral-like delivery system, Biochemistry, 43, 5618–5628, 2004. 

44. Lanza, G. M. et al., Novel paramagnetic contrast agents for molecular imaging and targeted drug 

delivery, Current Pharmaceutical Biotechnology, 5, 495–507, 2004. 

45. West, J. L. and Halas, N. J., Engineered nanomaterials for biophotonics applications: Improving 

sensing, imaging, and therapeutics, Annual Review of Biomedical Engineering, 5, 285–292, 2003. 

46. Stark, D. D. et al., Superparamagnetic iron-oxide—clinical-application as a contrast agent for MR 

imaging of the liver, Radiology, 168, 297–301, 1988. 

47. Choi, H., Choi, S. R., Zhou, R., Kung, H. F., and Chen, I. W., Iron oxide nanoparticles as magnetic 

resonance contrast agent for tumor imaging via folate receptor-targeted delivery, Academic 

Radiology, 11, 996–1004, 2004. 

48. Huh, Y. M. et al., In vivo magnetic resonance detection of cancer by using multifunctional magnetic 

nanocrystals, Journal of the American Chemical Society, 127, 12387–12391, 2005. 

49. Wang, S. J., Brechbiel, M., and Wiener, E. C., Characteristics of a new MRI contrast agent 

prepared from polypropyleneimine dendrimers, generation 2, Investigative Radiology, 38, 

662–668, 2003. 

50. Lanza, G. M. et al., Molecular imaging and targeted drug delivery with a novel, ligand-directed 

paramagnetic nanoparticle technology, Academic Radiology, 9, S330–S331, 2002. 

51. Loo, C. et al., Nanoshell-enabled photonics-based imaging and therapy of cancer, Technology in 

Cancer Research & Treatment, 3, 33–40, 2004. 

52. Sokolov, K. et al., Real-time vital optical imaging of precancer using anti-epidermal growth factor 

receptor antibodies conjugated to gold nanoparticles, Cancer Research, 63, 1999–2004, 2003. 

53. Loo, C., Lowery, A., Halas, N., West, J., and Drezek, R., Immunotargeted nanoshells for integrated 

cancer imaging and therapy, Nano Letters, 5, 709–711, 2005. 

54. Parak, W. J. et al., Biological applications of colloidal nanocrystals, Nanotechnology, 14, R15–R27, 

2003. 



55. Gao, X. H., Cui, Y. Y., Levenson, R. M., Chung, L. W. K., and Nie, S. M., In vivo cancer targeting 

and imaging with semiconductor quantum dots, Nature Biotechnology, 22, 969–976, 2004. 

56. Kim, S. et al., Near-infrared fluorescent type II quantum dots for sentinel lymph node mapping, 

Nature Biotechnology, 22, 93–97, 2004. 

57. Derfus, A. M., Chan, W. C. W., and Bhatia, S. N., Probing the cytotoxicity of semiconductor 

quantum dots, Nano Letters, 4, 11–18, 2004. 

58. Choi, Y. and Baker, J. R., Targeting cancer cells with DNA-assembled dendrimers—a mix and 

match strategy for cancer, Cell Cycle, 4, 669–671, 2005. 

59. Lu, Y., Yin, Y. D., Mayers, B. T., and Xia, Y. N., Modifying the surface properties of superparamagnetic 

iron oxide nanoparticles through a sol–gel approach, Nano Letters, 2, 183–186, 2002. 

60. Roy, I. et al., Ceramic-based nanoparticles entrapping water-insoluble photosensitizing anti-cancer 

drugs: A novel drug–carrier system for photodynamic therapy, Journal of the American Chemical 

Society, 125, 7860–7865, 2003. 

61. Lanza, G. M. et al., A novel site-targeted ultrasonic contrast agent with broad biomedical application, 

Circulation, 94, 3334–3340, 1996. 

62. Alkan-Onyuksel, H. et al., Development of inherently echogenic liposomes as an ultrasonic contrast 

agent, Journal of Pharmaceutical Sciences, 85, 486–490, 1996. 

63. Huang, S. L. et al., Improving ultrasound reflectivity and stability of echogenic liposomal dispersions 

for use as targeted ultrasound contrast agents, Journal of Pharmaceutical Sciences, 90, 1917–1926, 

2001. 

64. Rabin, O., Perez, J. M., Grimm, J., Wojtkiewicz, G., and Weissleder, R., An x-ray computed 

tomography imaging agent based on long-circulating bismuth sulphide nanoparticles, Nature 

Materials, 5, 118–122, 2006. 

65. Kelly, K. A. et al., Detection of vascular adhesion molecule-1 expression using a novel multimodal 

nanoparticle, Circulation Research, 96, 327–336, 2005. 

66. Morawski, A. M., Lanza, G. A., and Wickline, S. A., Targeted contrast agents for magnetic resonance 

imaging and ultrasound, Current Opinion in Biotechnology, 16, 89–92, 2005. 

67. Perez, J. M., Josephson, L., O’Loughlin, T., Hogemann, D., and Weissleder, R., Magnetic relaxation 

switches capable of sensing molecular interactions, Nature Biotechnology, 20, 816–820, 2002. 

68. Perez, J. M., Simeone, F. J., Saeki, Y., Josephson, L., and Weissleder, R., Viral-induced selfassembly 

of magnetic nanoparticles allows the detection of viral particles in biological media, 

Journal of the American Chemical Society, 125, 10192–10193, 2003. 

69. Perez, J. M., Simeone, F. J., Tsourkas, A., Josephson, L., and Weissleder, R., Peroxidase substrate 

nanosensors for MR imaging, Nano Letters, 4, 119–122, 2004. 

70. Georganopoulou, D. G. et al., Nanoparticle-based detection in cerebral spinal fluid of a soluble 

pathogenic biomarker for Alzheimer’s disease, Proceedings of the National Academy of Sciences 

of the United States of America, 102, 2273–2276, 2005. 

71. Mirkin, C. A., Letsinger, R. L., Mucic, R. C., and Storhoff, J. J., A DNA-based method for rationally 

assembling nanoparticles into macroscopic materials, Nature, 382, 607–609, 1996. 

72. Stevens, M. M., Flynn, N. T., Wang, C., Tirrell, D. A., and Langer, R., Coiled-coil peptide-based 

assembly of gold nanoparticles, Advanced Materials, 16, 915–918, 2004. 

73. Harris, T. J., von Maltzahn, G., Derfus, A. M., Ruoslahti, E., and Bhatia, S. N., Proteolytic actuation 

of nanoparticle self-assembly, Angewandte Chemie-International Edition, In press. 

74. Muller, R. H. and Keck, C. M., Challenges and solutions for the delivery of biotech drugs—a review 

of drug nanocrystal technology and lipid nanoparticles, Journal of Biotechnology, 113, 151–170, 

2004. 

75. Duncan, R., The dawning era of polymer therapeutics, Nature Reviews Drug Discovery, 2, 347–360, 

2003. 

76. Sessa, G. and Weissman, G., Phospholipid spherules (liposomes) as a model for biological 

membranes, Journal of Lipid Research, 9(3): 310–318, 1968. 

77. Gabizon, A. and Martin, F., Polyethylene glycol coated (pegylated) liposomal doxorubicin— 

rationale for use in solid tumours, Drugs, 54, 15–21, 1997. 


stability and membrane properties of paclitaxel-containing liposomes, Journal of Pharmaceutical 

Sciences, 90, 1091–1105, 2001. 


74 


79. Sapra, P. and Allen, T. M., Improved outcome when B-cell lymphoma is treated with combinations 

of immunoliposomal anti-cancer drugs targeted to both the CD19 and CD20 epitopes, Clinical 

Cancer Research, 10, 4893, 2004. 

80. Torchilin, V. P. et al., p-nitrophenylcarbonyl-PEG-PE-liposomes: Fast and simple attachment of 

specific ligands, including monoclonal antibodies, to distal ends of PEG chains via p-nitrophenylcarbonyl 

groups, Biochimica et Biophysica Acta-Biomembranes, 1511, 397–411, 2001. 

81. Park, J. W. et al., Tumor targeting using anti-her2 immunoliposomes, Journal of Controlled Release, 

74, 95–113, 2001. 

82. Miyamoto, M. et al., Preparation of gadolinium-containing emulsions stabilized with phosphatidylcholine-surfactant 

mixtures for neutron-capture therapy, Chemical & Pharmaceutical Bulletin, 47, 

203–208, 1999. 

83. Panyam, J. and Labhasetwar, V., Biodegradable nanoparticles for drug and gene delivery to cells and 

tissue, Advanced Drug Delivery Reviews, 55, 329–347, 2003. 

84. Nakanishi, T. et al., Development of the polymer micelle carrier system for doxorubicin, Journal of 

Controlled Release, 74, 295–302, 2001. 

85. Yokoyama, M. et al., Selective delivery of adiramycin to a solid tumor using a polymeric micelle 

carrier system, Journal of Drug Targeting, 7, 171–186, 1999. 

86. Torchilin, V. P., Structure and design of polymeric surfactant-based drug delivery systems, Journal 

of Controlled Release, 73, 137–172, 2001. 

87. Discher, B. M. et al., Polymersomes: Tough vesicles made from diblock copolymers, Science, 284, 

1143–1146, 1999. 

88. Ahmed, F. and Discher, D. E., Self-porating polymersomes of PEG-PLA and PEG-PCL: Hydrolysistriggered 

controlled release vesicles, Journal of Controlled Release, 96, 37–53, 2004. 

89. Xu, J. P., Ji, J., Chen, W. D., and Shen, J. C., Novel biomimetic polymersomes as polymer therapeutics 

for drug delivery, Journal of Controlled Release, 107, 502–512, 2005. 

90. Patri, A. K., Majoros, I. J., and Baker, J. R., Dendritic polymer macromolecular carriers for drug 

delivery, Current Opinion in Chemical Biology, 6, 466–471, 2002. 

91. Kojima, C., Kono, K., Maruyama, K., and Takagishi, T., Synthesis of polyamidoamine dendrimers 

having poly(ethylene glycol) grafts and their ability to encapsulate anti-cancer drugs, Bioconjugate 

Chemistry, 11, 910–917, 2000. 

92. Patri, A. K. et al., Synthesis and in vitro testing of J591 antibody-dendrimer conjugates for targeted 

prostate cancer therapy, Bioconjugate Chemistry, 15, 1174–1181, 2004. 

93. Tripathi, P. K. et al., Dendrimer grafts for delivery of 5-flurouracil, Pharmazie, 57, 261–264, 2002. 

94. Quintana, A. et al., Design and function of a dendrimer-based therapeutic nanodevice targeted to 

tumor cells through the folate receptor, Pharmaceutical Research, 19, 1310–1316, 2002. 

95. Vinogradov, S. V., Batrakova, E. V., and Kabanov, A. V., Nanogels for oligonucleotide delivery to 

the brain, Bioconjugate Chemistry, 15, 50–60, 2004. 

96. Jordan, A., Scholz, R., Wust, P., Fahling, H., and Felix, R., Magnetic fluid hyperthermia (MFH): 

Cancer treatment with AC magnetic field induced excitation of biocompatible superparamagnetic 

nanoparticles, Journal of Magnetism and Magnetic Materials, 201, 413–419, 1999. 

97. Pankhurst, Q. A., Connolly, J., Jones, S. K., and Dobson, J., Applications of magnetic nanoparticles 

in biomedicine, Journal of Physics D—Applied Physics, 36, R167–R181, 2003. 

98. Derfus, A. M. and Bhatia, S. N., Unpublished data. 

99. Rabin, Y., Is intracellular hyperthermia superior to extracellular hyperthermia in the thermal sense? 

International Journal of Hyperthermia, 18, 194–202, 2002. 

100. Hirsch, L. R. et al., Nanoshell-mediated near-infrared thermal therapy of tumors under magnetic 

resonance guidance, Proceedings of the National Academy of Sciences of the United States of 

America, 100, 13549–13554, 2003. 

101. O’Neal, D. P., Hirsch, L. R., Halas, N. J., Payne, J. D., and West, J. L., Photo-thermal tumor ablation 

in mice using near infrared-absorbing nanoparticles, Cancer Letters, 209, 171–176, 2004. 

102. Plank, C. et al., The magnetofection method: Using magnetic force to enhance gene delivery, 

Biological Chemistry, 384, 737–747, 2003. 

103. Nishiyama, N. et al., Light-harvesting ionic dendrimer porphyrins as new photosensitizers for 

photodynamic therapy, Bioconjugate Chemistry, 14, 58–66, 2003. 



104. Konan, Y. N., Gurny, R., and Allemann, E., State of the art in the delivery of photosensitizers for 

photodynamic therapy, Journal of Photochemistry and Photobiology B—Biology, 66, 89–106, 2002. 

105. Kopelman, R. et al., Multifunctional nanoparticle platforms for in vivo MRI enhancement and 

photodynamic therapy of a rat brain cancer, Journal of Magnetism and Magnetic Materials, 293, 

404–410, 2005. 

106. Crowder, K. C. et al., Sonic activation of molecularly-targeted nanoparticles accelerates transmembrane 

lipid delivery to cancer cells through contact-mediated mechanisms: Implications for 

enhanced local drug delivery, Ultrasound in Medicine and Biology, 31, 1693–1700, 2005. 

107. Gao, Z. G., Fain, H. D., and Rapoport, N., Controlled and targeted tumor chemotherapy by micellarencapsulated 

drug and ultrasound, Journal of Controlled Release, 102, 203–222, 2005. 

108. Hainfeld, J. F., Slatkin, D. N., and Smilowitz, H. M., The use of gold nanoparticles to enhance 

radiotherapy in mice, Physics in Medicine and Biology, 49, N309–N315, 2004. 

109. Haag, R., Supramolecular drug-delivery systems based on polymeric core-shell architectures, Angewandte 

Chemie-International Edition, 43, 278–282, 2004. 

110. Murthy, N., Campbell, J., Fausto, N., Hoffman, A. S., and Stayton, P. S., Design and synthesis of 

pH-responsive polymeric carriers that target uptake and enhance the intracellular delivery of oligonucleotides, 

Journal of Controlled Release, 89, 365–374, 2003. 

111. Forster, S., Abetz, V., and Muller, A. H. E., Polyelectrolyte block copolymer micelles, Polyelectrolytes 

with Defined Molecular Architecture Ii, 166, 173–210, 2004. 


6 

CONTENTS 

Neutron Capture Therapy of 

Cancer: Nanoparticles and High 

Molecular Weight Boron 

Delivery Agents 

Gong Wu, Rolf F. Barth, Weilian Yang, Robert J. Lee, 

Werner Tjarks, Marina V. Backer, and Joseph M. Backer 

6.1 Introduction ......................................................................................................................... 78 

6.2 General Requirements for Boron Delivery Agents ............................................................ 78 

6.3 Low-Molecular-Weight Delivery Agents ........................................................................... 79 

6.4 High-Molecular-Weight Boron Delivery Agents ............................................................... 79 

6.5 Dendrimer-Related Delivery Agents .................................................................................. 80 

6.5.1 Properties of Dendrimers ........................................................................................ 80 

6.5.2 Boronated Dendrimers Linked to Monoclonal Antibodies .................................... 80 

6.5.2.1 Boron Clusters Directly Linked to mAb ................................................. 80 

6.5.2.2 Attachment of Boronated Dendrimers to mAb ....................................... 81 

6.5.3 Boronated Dendrimers Delivered by Receptor Ligands ........................................ 81 

6.5.3.1 Epidermal Growth Factors (EGF) ........................................................... 81 

6.5.3.2 Folate Receptor-Targeting Agents........................................................... 83 

6.5.3.3 Vascular Endothelial Growth Factor (VEGF) ......................................... 83 

6.5.4 Other Boronated Dendrimers .................................................................................. 85 

6.6 Liposomes as Boron Delivery Agents ................................................................................ 86 

6.6.1 Overview of Liposomes .......................................................................................... 86 

6.6.2 Boron Delivery by Nontargeted Liposomes ........................................................... 86 

6.6.2.1 Liposomal Encapsulation of Sodium Borocaptate and 

Boronophenylalanine ............................................................................... 86 

6.6.2.2 Liposomal Encapsulation of Other Boranes and Carboranes ................. 87 

6.6.3 Boron Delivery by Targeted Liposomes ................................................................ 89 

6.6.3.1 Immunoliposomes .................................................................................... 89 

6.6.3.2 Folate Receptor-Targeted Liposomes ...................................................... 90 

6.6.3.3 EGFR-Targeted Liposomes ..................................................................... 91 

6.7 Boron Delivery by Dextrans ............................................................................................... 92 

Experimental studies described in this report were supported by N.I.H. grants 1 R01 CA098945 (R.F. Barth), 1 R01 

CA79758 (W. Tjarks), and Department of Energy grants DE-FG02-98ER62595 (R.F. Barth) and DE-FG-2-02FR83520 

(J.M. Backer). 


77

78 

6.8 Other Macromolecules Used for Delivering Boron Compounds....................................... 93 

6.9 Delivery of Boron-Containing Macromolecules to Brain Tumors .................................... 93 

6.9.1 General Considerations ........................................................................................... 93 

6.9.2 Drug-Transport Vectors .......................................................................................... 93 

6.9.3 Direct Intracerebral Delivery .................................................................................. 94 

6.9.4 Convection-Enhanced Delivery .............................................................................. 94 

6.10 Clinical Considerations and Conclusions ......................................................................... 95 

6.11 Summary ........................................................................................................................... 96 

Acknowledgments ........................................................................................................................ 96 

References .................................................................................................................................... 96 


After decades of intensive research, high-grade gliomas, and specifically glioblastoma multiforme 

(GBM), are still extremely resistant to all current forms of therapy including surgery, 

chemotherapy, radiotherapy, immunotherapy and gene therapy. 1–5 The five-year survival rate of 

patients diagnosed with GBM in the United States is less than a few percent 6,7 despite aggressive 

treatment using combinations of therapeutic modalities. This is due to the infiltration of malignant 

cells beyond the margins of resection and their spread into both gray and white matter by the time of 

surgical resection. 8,9 High-grade gliomas are histologically complex and heterogeneous in their 

cellular composition. Recent molecular genetic studies of gliomas have shown how complex the 

development of these tumors is. 10 Glioma cells and their neoplastic precursors have biologic 

properties that allow them to evade a tumor associated host immune response, 11 and biochemical 

properties that allow them to invade the unique extracellular environment of the brain. 12,13 Consequently, 

high-grade supratentorial gliomas must be regarded as a whole-brain disease. 14 The 

inability of chemo- and radiotherapy to cure patients with high-grade gliomas is due to their 

failure to eradicate micro-invasive tumor cells within the brain. To successfully treat these 

tumors, strategies must be developed that can selectively target malignant cells with little or no 

effect on normal cells and tissues adjacent to the tumor. Boron neutron capture therapy (BNCT) is 

based on the nuclear capture and fission reactions that occur when non radioactive boron-10 is 

irradiated with low-energy thermal neutrons to yield high-linear-energy-transfer (LET) alpha 

particles ( 4 He) and recoiling lithium-7 ( 7 Li) nuclei. For BNCT to be successful, a sufficient 

number of 10 B atoms (approximately 10 9 atoms/cell) must be delivered selectively to the tumor 

and enough thermal neutrons must be absorbed by them to sustain a lethal 10 B(n,a) 7 Li-capture 

reaction. The destructive effects of these high-energy particles are limited to boron-containing cells. 

BNCT primarily has been used to treat high-grade gliomas, 15,16 and either cutaneous primaries 17 or 

cerebral metastases of melanoma. 18 More recently, it also has been used to treat patients with headand-neck 

19,20 and metastatic liver cancer. 21,22 BNCT is a biologically rather than physically 

targeted type of radiation treatment. If sufficient amounts of 10 B and thermal neutrons can be 

delivered to the target volume, the potential exists to destroy tumor cells dispersed in the normal 

tissue parenchyma. Readers interested in more in depth coverage of other topics related to 

BNCT are referred to several recent reviews 23–25 and monographs. 15,24 This review will focus 

on boron-containing macromolecules and nanovehicles as boron delivery agents. 

6.2 GENERAL REQUIREMENTS FOR BORON DELIVERY AGENTS 


A successful boron delivery agent should have (1) no or minimal systemic toxicity with rapid 

clearance from blood and normal tissues, (2) high tumor (approximately 20 mg 10 B/g) and low 


Neutron Capture Therapy of Cancer 79 

normal tissue uptake, (3) high tumor to brain (T:Br) and tumor to blood (T:Bl) concentration ratios 

(greater than 3–4:1), and (4) persistence in the tumor for a sufficient period of time to carry out 

BNCT. At this time, no single boron delivery agent fulfills all of these criteria. However, as a result 

of new synthetic techniques and increased knowledge of the biological and biochemical requirements 

for an effective agent, multiple new boron agents have emerged and these are described in a 

special issue of Anti-Cancer Agents in Medical Chemistry (6, 2, 2006). The major challenge in their 

development has been the requirement for specific tumor targeting to achieve boron concentrations 

sufficient to deliver therapeutic doses of radiation to the tumor with minimal normal-tissue toxicity. 

The selective destruction of GBM cells in the presence of normal cells represents an even greater 

challenge compared to malignancies at other anatomic sites. 

6.3 LOW-MOLECULAR-WEIGHT DELIVERY AGENTS 

In the 1950s and early 1960s clinical trials of BNCT were carried out using boric acid and some of 

its derivatives as delivery agents. These simple chemical compounds had poor tumor retention, 

attained low T:Br ratios and were nonselective. 26,27 Among the hundreds of low-molecular-weight 

(LMW) boron-containing compounds that were synthesized, two appeared to be promising. One, 

based on arylboronic acids, 28 was (L)-4-dihydroxy-borylphenylalanine, referred to as boronophenylalanine 

or BPA (Figure 6.1, 1). The second, a polyhedral borane anion, was sodium 

mercaptoundecahydro-closo-dodecaborate, 29 more commonly known as sodium borocaptate or 

BSH (2). These two compounds persisted longer in animal tumors compared with related 

molecules, attained T:Br and T:Bl boron ratios greater than 1 and had low toxicity. 10 B-enriched 

BPA, complexed with fructose to improve its water solubility, and BSH, have been used clinically 

for BNCT of brain, as well as extracranial tumors. Although their selective accumulation in tumors 

is not ideal, the safety of these two drugs following intravenous (i.v.) administration has been well 

established. 30,31 

6.4 HIGH-MOLECULAR-WEIGHT BORON DELIVERY AGENTS 

High-molecular-weight (HMW) delivery agents usually contain a stable boron group or cluster 

linked via a hydrolytically stable bond to a tumor-targeting moiety, such as monoclonal antibodies 

(mAbs) or LMW receptor-targeting ligands. Examples of these include epidermal growth factor 

(HO) 2 B 

OCN 

1 

3 

H 

CO 2 – 

NH 3 + 

N(CH 3) 3Na 

FIGURE 6.1 Structures of two compounds used clinically for boron neutron capture therapy (BNCT), 4-dihydroxyborylphenylalanine 

or boronophenylalanine (BPA, 1), and disodium undecahydro-closo-dodecaborate or 

sodium borocaptate (BSH, 2) and the isocyanato polyhedral borane (3), which has been used to heavily 

boronate dendrimers. 


Na 2 

SH 

2 

BH 

B

80 

(EGF) and the mAb cetuximab (IMC-C225) to target the EGF receptor (EGFR) or its mutant 

isoform EGFRvIII, which are overexpressed in a variety of malignant tumors including gliomas, 

and squamous-cell carcinomas of the head and neck. 32 Agents that are to be administered systemically 

should be water soluble (WS), but lipophilicity is important to cross the blood–brain barrier 

(BBB) and diffuse within the brain and the tumor. There should be a favorable differential in boron 

concentrations between tumor and normal brain, thereby enhancing their tumor specificity. Their 

amphiphilic character is not as crucial for LMW agents that target-specific biological transport 

systems and/or are incorporated into nanovehicles such as liposomes. Molecular weight also is an 

important factor because it determines the rate of diffusion both within the brain and the tumor. 

Detailed reviews of the state-of-the-art of compound development for BNCT have been 

published. 33,34 In this review, we will focus on boron containing macromolecules and liposomes 

as delivery agents for BNCT, and how they can be most effectively administered. 

6.5 DENDRIMER-RELATED DELIVERY AGENTS 

6.5.1 PROPERTIES OF DENDRIMERS 

Dendrimers are synthetic polymers with a well-defined globular structure. As shown in Figure 6.2, 

they are composed of a core molecule, repeat units that have three or more functionalities, and 

reactive surface groups. 35,36 Two techniques have been used to synthesize these macromolecules: 

divergent growth outwards from the core, 37 or convergent growth from the terminal groups inwards 

towards the core. 36,38 Regular and repeated branching at each monomer group gives rise to a 

symmetric structure and pattern to the entire globular dendrimers. Dendrimers are an attractive 

platform for macromolecular imaging and gene delivery because of their low cytotoxicity and their 

multiple types of reactive terminal groups. 36,39–44 

6.5.2 BORONATED DENDRIMERS LINKED TO MONOCLONAL ANTIBODIES 

6.5.2.1 Boron Clusters Directly Linked to mAb 

Monoclonal antibodies (mAb) have been attractive targeting agents for delivering radionuclides, 45 

drugs, 46–50 toxins, 51 and boron to tumors. 52–55 Prior to the introduction of dendrimers as boron 

carriers, boron compounds were directly attached to mAbs. 53,54 It has been calculated that 

Terminal groups 

R HN 

Linkers 

= Repeating units 

R = mAb, EGF, Folic acid or VEGF 

H 

N H 

N – 

N(CH3 ) 3 

FIGURE 6.2 Structure of a boronated PAMAM dendrimer that has been linked to targeting moieties. 

PAMAM dendrimers consist of a core, repeating polyamido amino units, and reactive terminal groups. 

Each successively higher generation of PAMAM dendrimer has a geometrically incremental number of 

terminal groups. Boronated dendrimers have been prepared by reacting them with water-soluble isocyanato 

polyhedral boranes and subsequently attaching them to targeting moieties by means of heterobifunctional 

linkers. 



O 

BH 

B


approximately 10 9 10 B atoms per cell (approximately 20 mg/g tumor) must be delivered to kill 

tumor cells. 55,56 Based on the assumption of 10 6 antigenic receptor sites per cell, approximately 50– 

100 boron cage structures of carboranes, or polyhedral borane anions and their derivatives must be 

linked to each mAb molecule to deliver the required amount of boron for NCT. The attachment of 

such a large number of boron cages to a mAb may result in precipitation of the bioconjugate or a 

loss of its immunological activity. Solubility can be improved by inserting a water-soluble gluconamide 

group into the protein-binding boron cage compounds, thereby enhancing their water 

solubility. 57 This modification makes it possible to incorporate up to 1100 boron atoms into a 

human gamma globulin (HGG) molecule without any precipitation. Other approaches to 

enhance solubility include the use of negatively charged carboranes 58 or polyhedral borane 

anions, 59 as well as the insertion of carbohydrate groups. 60,61 A major limitation of using an 

agent containing a single boron cage is that a large number of sites must be modified to deliver 

10 3 boron atoms per molecule of antibody and this can reduce its immunoreactivity activity. Alam 

et al. showed that attachment of an average of 1300 boron atoms to mAb 17-1A, which is directed 

against human colorectal carcinoma cells, resulted in a 90% loss of its immunoreactivity. 62 

6.5.2.2 Attachment of Boronated Dendrimers to mAb 

Dendrimers are one of the most attractive polymers that have been used as boron carriers due to 

their well-defined structure and large number of reactive terminal groups. Depending on the antigen 

site density, approximately 1000 boron atoms need to be attached per molecule of dendrimer and 

subsequently linked to the mAb. In our first study, second- and fourth-generation polyamido amino 

(PAMAM or “starburst”) dendrimers, which have 12 and 48 reactive terminal amino groups, 

respectively, were reacted with the water-soluble isocyanato polyhedral borane [Na(CH3)3NB10 

H8NCO] (3). 63,64 The boronated dendrimer then was linked to the mAb IB16-6, which is directed 

against the murine B16 melanoma, by means of two heterobifunctional linkers, m-maleimidobenzoyl-N-hydroxysulfosuccinimide 

ester (sulfo-MBS) and N-succinimidyl 3-(2-pyridyldithio) 

propionate (SPDP). 63,65 However, following i.v. administration, large amounts of the bioconjugate 

accumulated in the liver, and spleen and it was concluded that random conjugation of boronated 

dendrimers to a mAb could alter its binding affinity and biodistribution. To minimize the loss of 

mAb reactivity, a fifth-generation PAMAM dendrimer was boronated with the same polyhedral 

borane anion, and more recently it was site-specifically linked to the anti-EGFR mAb cutuximab 

(or IMC-C225) or the EGFRvIII-specific mAb L8A4 (Figure 6.3). Cetuximab was linked via 

glycosidic moieties in the Fc region by means of two heterobifunctional reagents, SPDP and 

N-(k-maleimidoundecanoic acid) hydrazide (KMUH). 66,67 The resulting bioconjugate, designated 

C225-G5-B 1100, contained approximately 1100 boron atoms per molecule of cetuximab and 

retained its aqueous solubility in 10% DMSO and its in vitro and in vivo immunoreactivity. As 

determined by a competitive binding assay, there was a less than 1 log unit decrease in affinity for 

EGFR (C) glioma cell line F98EGFR, compared to that of unmodified cetuximab. 66 In vivo biodistribution 

studies, carried out 24 h after intratumoral (i.t) administration of the bioconjugate, 

demonstrated that 92.3 mg/g of boron was retained in rats bearing F98EGFR gliomas, compared to 

36.5 mg/g in EGFR (K) F98 parental tumors and 6.7 mg/g in normal brain. 67 

6.5.3 BORONATED DENDRIMERS DELIVERED BY RECEPTOR LIGANDS 

6.5.3.1 Epidermal Growth Factors (EGF) 

Due to its increased expression in a variety of tumors, including high-grade gliomas, and its low or 

undetectable expression in normal brain, EGFR is an attractive target for cancer therapy. 68–70 

As described above, targeting of EGFR has been carried out using either mAbs or alternatively, 

as described in this section, EGF, which is a single-chain, 53-mer heat and acid stable polypeptide. 


82 

HO 

O 

HO 

O 

O 

O 

O O 

O 

H 

N S 

O 

H 

N H N 

O 

O 

OH 

OH 

BH 

B 

Cetuximab (C225) 

NalO 4 

C225-CHO 

N(CH 3) 3 – 

KMUH-G5-B 1100 

O O 

O O 

N 

O 

(CH2 ) 10 N 

H 

N 

O 

O OH 

C225-G5-B 1100 

NH 2 

NH 2 

5th generation PAMAM dendrimer 

O 

S 

S 

NH 2 

H 

N 

H 

N 

N 

O 

SPDP-G5 

H 

N SH 

O 

HS-G5-B1100 O 

H 

N 

O 

N 

O 

O 

NH2 (CH2 ) 10 N 

H 

S 

KMUH-G5-B 1100 

S S N 

Na(CH 3 ) 3 NB 10 H 8 NCO 

O 

SPDP-G5-B 1100 

DTT 

S S N 

O 

N 

O 

O 

N 

O 

(CH 2 ) 10 

FIGURE 6.3 Conjugation scheme for linkage of a boron-containing dendrimer to the monoclonal antibody, 

C225 (cetuximab), which is directed against EGFR. (From Wu, G. et al., Bioconjug. Chem., 15, 185, 2004; 

Barth, R. F. et al., Appl. Radiat. Isot. 61, 899, 2004. With permission.) 



O 

N 

H 

NH 2


It binds to a transmembrane glycoprotein with tyrosine kinase activity, which triggers dimerization 

and internalization. 71,72 Because the EGF boron bioconjugates have a much smaller MW than mAb 

conjugates, they should be capable of more rapid and effective tumor targeting than has been 

observed with mAbs. 67,73 

The procedure used to conjugate EGF to a boronated dendrimer was slightly different from that 

used to boronate mAbs. A fourth-generation PAMAM dendrimer was reacted with the isocyanato 

polyhedral borane Na(CH3)3NB10H8NCO. Next, reactive thiol groups were introduced into the 

boronated dendrimer using SPDP, and EGF was derivatized with sulfo-MBS. The reaction of thiol 

groups of the derivatized, boronated starburst dendrimer (BSD) with maleimide groups produced a 

stable BSD–EGF bioconjugate, which contained approximately 960 atoms of boron per molecule of 

EGF. 74 The BSD–EGF initially was bound to the cell surface membrane and then was endocytosed, 

which resulted in accumulation of boron in lysosomes. 74 Subsequently, in vitro and in vivo studies 

were carried out to evaluate the potential efficacy of the bioconjugate as a boron delivery agent for 

BNCT. 73 As will be described in more detail later on in this review, therapy studies demonstrated 

that F98EGFR-glioma-bearing rats that received either boronated EGF or mAb by either direct i.t. 

injection or convection enhanced delivery into the brain had a longer mean survival time (MST) 

than animals bearing F98 parental tumors following BNCT. 75–77 

6.5.3.2 Folate Receptor-Targeting Agents 

Folate receptor (FR) is overexpressed on a variety of human cancers, including those originating in 

ovary, lung, breast, endometrium and kidney. 78–80 Folic acid (FA) is a vitamin that is transported 

into cells via FR mediated endocytosis. It has been well documented that the attachment of FA via 

its g-carboxylic function to other molecules does not alter its endocytosis by FR-expressing cells. 81 

FR targeting has been used successfully to deliver protein toxins, chemotherapeutic, radioimaging, 

therapeutic and MRI contrast agents, 82 liposomes, 83 gene transfer vectors, 84 antisense oligonucleotides, 

85 ribozymes, and immunotherapeutic agents to FR-positive cancers. 86 To deliver boron 

compounds, FA was conjugated to heavily boronated third generation PAMAM dendrimers 

containing polyethylene glycol (PEG). 87 PEG was introduced into the bioconjugate to reduce its 

uptake by the reticuloendothelial system (RES), and more specifically, the liver and spleen. It was 

observed that folate linked to third generation PAMAM dendrimers containing 12–15 decaborate 

clusters and 1–1.5 PEG2000 units had the lowest hepatic uptake in C57Bl/6 mice (7.2–7.7% injected 

dose [I.D.]/g liver). In vitro studies using FR (C) KB cells demonstrated receptor-dependent uptake 

of the bioconjugate. Biodistribution studies with this conjugate, carried out in C57Bl/6 mice 

bearing subcutaneous (s.c.) implants of the FR (C) murine sarcoma 24JK-FBP, demonstrated 

selective tumor uptake (6.0% I.D./g tumor), but there was high hepatic (38.8% I.D./g) and renal 

(62.8% I.D./g) uptake. 87 

6.5.3.3 Vascular Endothelial Growth Factor (VEGF) 

There is preclinical and clinical evidence indicating that angiogenesis plays a major role in the 

growth and dissemination of malignant tumors. 88,89 Inhibition of angiogenesis has yielded 

promising results in a number of experimental animal tumor models. 90,91 The most important 

molecular targets have been vascular endothelial growth factor (VEGF) and its receptor 

(VEGFR). 92,93 We recently have constructed a human VEGF121 isoform fused to a novel 15-aa 

cysteine-containing tag designed for site-specific conjugation of therapeutic and diagnostic 

agents 94 (Figure 6.4). A boronated fifth-generation PAMAM dendrimer (BD), tagged with a 

near-infrared (IR) fluorescent dye Cy5, was conjugated using the heterobifunctional reagent 

sulfo-LC–SPDP to BD/Cy5 via reactive SH-groups generated in the VEGF fusion protein by 

mild dithiothreitol (DTT) treatment. The bioconjugate, designated VEGF–BD/Cy5, contained 

1050–1100 boron atoms per dimeric VEGF molecule. VEGF–BD/Cy5 retained the in vitro 


84 

H 2N 

Cy5 

Cy5 

A 

B 

G5 

G5 

G5 

Cy5 

NH 2 

NH 2 

G5 

NH 2 

SPDP 

SPDP 

Cy5 

S 

S 

G5 

SH 

SH 

D 

S 

S 

VEGF-BD-Cy5 

H 

N H N 

functional activity of VEGF121, which was similar to that of native VEGF. Tagging the bioconjugate 

with Cy5 dye permitted near-IR fluorescence imaging of its in vitro and in vivo uptake. 

In vitro uptake was determined by incubating VEGF–BD/Cy5 with VEGFR-2 overexpressing 

PAE/KDR cells and in vivo uptake was evaluated in mice bearing s.c. tumor implants. In vitro, 

HS 

G5 

S 

S 

O 

SH 

Cy5 

G5 

C 


BH 

B 

Cy5 

SH 

N(CH 3 ) 3 – 

FIGURE 6.4 Preparation of a vascular endothelial growth factor (VEGF) receptor-targeting nanovehicle 

(VEGF–BD/Cy5) (A). A fifth-generation (G5) PAMAM dendrimer initially is reacted with SPDP and the 

near-IR dye, Cy5-NHS, (A) dissolved in dimethylformamide/methanol mixture for 1 h. Following this, it is 

reacted with the polyisocynato polyhedral borane, Na(CH 3) 3–NB 10H 8NCO, in a carbonate buffer (pH 9.0)/9% 

acetone mixture for 24 h (B). Then a 1.5-fold molar excess of DTT is added to the VEGF monomer in a 

solution containing 20 mmol/L NaOAc (pH 6.5), 0.5 mol/L urea and 0.5 mol/L NDSB-221 and incubated at 

48C for 16 h (C). Finally, the boronated dendrimer is incubated with the targeting vehicle for 15 min to produce 

the VEGF–BD/Cy5 (D). (From Backer, M. V. et al., Mol. Cancer Ther., 4, 1423, 2005.) 


SH


the bioconjugate localized in the perinuclear region and in vivo it primarily localized in regions of 

active angiogenesis. Furthermore, depletion of VEGFR-2 overexpressing cells from the tumor 

vasculature with VEGF-toxin fusion protein significantly decreased bioconjugate uptake, indicating 

that these cells were the primary targets of VEGF–BD/Cy5. These studies demonstrated that 

VEGF–BD/Cy5 potentially could be used as a diagnostic agent. 94 Further studies are planned to 

evaluate its therapeutic efficacy using the F98 rat glioma model, which we have used extensively 

to evaluate both LMW and HMW boron delivery agents. 95 

6.5.4 OTHER BORONATED DENDRIMERS 

An alternative method to deliver boron compounds by means of dendrimers is to incorporate 

carborane cages within the dendrimer (Figure 6.5). Matthew et al. have reported that 4, 8, or 16 

carboranes could be inserted into an aliphatic polyester dendrimer by means of a highly effective 

HO 

HO 

HO 

HO 

HO 

O 

O 

O 

O 

O 

O 

HO 

O 

OH 

O 

O 

O O 

O O 

O 

O O 

O 

O 

O O 

O 

O 

O 

OH 

O 

O 

O 

O 

O 

O 

O 

O 

O 

O 

O 

O O O 

FIGURE 6.5 Structure of a barborane-containing aliphatic polyester dendrimer. Carborane cages were incorporated 

into the interior of the dendrimer structure and the peripheral hydrophilic groups improved water 

solubility and were available for modification. 


HO 

O 

O O 

O 

O 

O O 

O 

O 

O 

O 

O 

HO 

O 

OH 

O 

O 

O 

O 

O 

O 

OH 

OH 

BH 

C 

OH 

OH 

OH

86 

synthon, a bifunctional carborane derivative with an acid group and a benzyl ether protected 

alcohol. 96 The procedure employed a divergent synthesis with high yield. The resulting polyhydroxylated 

dendrimer was WS with a minimum ratio of eight hydroxyl groups per carborane cage. 

Carborane containing dendrimers potentially could be used as boron delivery agents for BNCT 

because it is possible to control the number of carborane moieties and overall solubility. 

6.6 LIPOSOMES AS BORON DELIVERY AGENTS 

6.6.1 OVERVIEW OF LIPOSOMES 

Liposomes are biodegradable, nontoxic vesicles that have been used to deliver both hydrophilic and 

hydrophobic agents (Figure 6.6). 97 Both classical and PEGylated (“stealth”) liposomes can increase 

the amounts of anti-cancer drugs that can be delivered to solid tumors by passive targeting. Rapidly 

growing solid tumors have increased permeability to nanoparticles due to increased capillary pore 

size. These can range from 100 to 800 nm. In comparison, endothelial pore size of normal tissues, 

which are impermeable to liposomes, can range in size from 60 to 80 nm. In addition, tumors lack 

efficient lymphatic drainage, and consequently, clearance of extravasated liposomes is slow. 98 

Modification of the liposomal surface by PEGylation or attachment of antibodies or receptor 

ligands, will improve their selective targeting and increase their circulation time. 98 

6.6.2 BORON DELIVERY BY NONTARGETED LIPOSOMES 

6.6.2.1 Liposomal Encapsulation of Sodium Borocaptate and Boronophenylalanine 

Liposomes have been extensively evaluated as nanovehicles for the delivery of boron compounds 

for NCT. 99,100 In vitro and in vivo studies have demonstrated that they can effectively and selectively 

deliver large quantities of boron to tumors and that the compounds delivered by liposomes 

have a longer tumor retention time. BPA is an amino acid analogue that is preferentially taken up by 

cells with increased metabolic activity, such as tumor cells of varying histopathologic types 

Bilayer memberane Internal aqueous space 


5nm 

Hydrocarbon tails 

Polar head groups 

mAb or ligand 

FIGURE 6.6 Schematic diagram of the structure of a liposome that has an aqueous core and a lipid bilayer 

membrane. The latter is composed of polar head groups with hydrocarbon tails. The liposomal surface can be 

modified by PEGylation to prolong its circulation time and linked to either a mAb or a ligand for targeting. 


PEG


including melanomas, 31,101 gliomas 102 , and squamous cell carcinomas. 103,104 Because of its low 

aqueous solubility, BPA has been used as a fructose complex, which has permitted it to be administered 

i.v. rather than orally. 105,106 Following i.v. administration of BPA, which had been 

incorporated into conventional liposomes, there was rapid elimination by the RES with very low 

blood boron concentrations at 3 h. In contrast, if BPA was incorporated into liposomes composed of 

distearoyl phosphoethanolamine (DSPE)–PEG, therapeutically effective tumor boron concentrations 

(greater than 20 mg/g) were seen at 3 h and at 6 h, indicating that PEG-liposomes had 

evaded the RES. 107 In addition, BPA has been incorporated into the lipid bilayer of liposomes, 

composed of positively charged lipid 1,2-dioleoyl-3-trimethylammonium-propane (DOTAP) and 

the zwitterionic lipid, 1,2-dioleoyl-sn-glycerol-3-phosphoethanolamine (DOPE). 108 Cationic liposomes 

have been widely used as carriers of biomolecules that specifically target the cell nucleus, 109 

which would be advantageous for BNCT. Another clinically used drug, BSH, has been incorporated 

into liposomes composed of dipalmitoylphosphatidylcholine (DPPC)/Chol in a 1:1 molar ratio with 

and without PEG stabilization. 110 The average diameter of liposomes containing BSH was in the 

range of 100–110 nm. Both types of liposomes resulted in a significant improvement in their 

circulation time compared to that of free BSH. At 24 h following i.v. injection of PEG-liposomes, 

19% of the I.D. of boron was in the blood compared to 7% following formulation of BSH in 

conventional liposomes. The mean percent uptake by the liver and spleen was not significantly 

different for the two types of liposomes. However, the blood to RES ratios were higher for PEGliposomes 

at all time points indicating that a higher fraction of the injected dose (I.D.) of BSH was 

still in the blood. Ji et al. have reported that there were no significant differences in the in vitro 

uptake by 9L gliosarcoma cells of free BSH versus a liposomal formulation after 16 h incubation. 

However, cellular persistence was increased at 12 and 24 h for BSH-loaded liposomes. 111 BSH also 

has been incorporated into transferrin (TF) conjugated PEG liposomes (TF–PEG liposomes), 112 

which then were taken up by cells via TF receptor-mediated endocytosis. Intravenous administration 

of this formulation increased boron retention at the tumor site compared with PEG 

liposomes, bare liposomes or free BSH and suppressed tumor growth following BNCT. These 

results suggest that TF targeted liposomes might be useful as intracellular targeting vehicles. 

6.6.2.2 Liposomal Encapsulation of Other Boranes and Carboranes 

Polyhedral boranes 34 and carboranes 113,114 are another class of boron compounds that have been 

used for NCT. They contain multiple boron atoms per molecule, are resistant to metabolic 

degradation, and are lipophilic, thereby permitting easier penetration of the tumor cell 

membrane. 113 In addition to BSH, carboranylpropylamine (CPA, Figure 6.7, 4) 115 has been incorporated 

into conventional and PEGylated liposomes by active loading, using a transmembrane pH 

gradient. 115,116 Although as many as 13,000 molecules of CPA were loaded into liposomes having a 

mean average diameter of 100 nm, there was no in vitro toxicity to both the glioblastoma cell line 

SK-MG-1 and normal human peripheral blood lymphocytes. Borane anions, such as B10H 2K 

10 , 

B12H11SH 2K ,B20H17OH 4K , and B20H 3K 

19 , and the normal form and photoisomer of B20H 2K 

18 , also 

have been encapsulated into small unilamellar vesicles with mean diameters of 70 nm or less. 117–120 

These were composed of the pure synthetic phospholipids, distearoyl phosphatidylcholine, and 

cholesterol. 117 Although encapsulation efficiencies were only 2–3%, following i.v. injection, 

these liposomes were selectively delivered to the tumors in mice bearing the EMT6 mammary 

carcinoma and had attained boron concentrations of greater than 15 mg boron/g, with a T:Bl ratio of 

greater than 3. Two isomers of B 20H 2K 

18 attained boron concentrations of 13.6 and 13.9 mg/g with 

T:Bl ratios of 3.3 and 12, respectively, at 48 h following administration. High boron retention and 

prolongation of their circulation time was observed due to the interaction with intracellular components 

after it had been released from liposomes within tumor cells. 117 

To examine the effect of charge and substitution on the retention of boranes, two isomers 

[B 20H 17NH 3] 3K and [1-(2 0 -B 10H 9)-2-NH 3B 10H 8] 3K (5) were prepared from the polyhedral 


88 

O O 

H 

4 

6 

8 

NH 3 + Cl – 

2– 

(CH 2 ) 15 CH 3 

– 

– + 

ClH3N(CH2 ) 3 

– + 

ClH3N(CH2 ) 4 

+Cl 

HN(CH2 ) 3 

– 

10 11 


borane anion [n-B 20H 18] 2K (6). 118,119 The sodium salts of these two isomers had been encapsulated 

within small unilamellar liposomes, composed of distearoyl phosphatidylcholine/cholesterol at a 

1:1 ratio. Both isomers of [B 20H 17NH 3] 3K had excellent tumor uptake and selectivity in EMT 6 

tumor bearing mice, even at very low I.D.s, and this resulted in peak tumor boron concentrations of 

30–40 mg B/g and a T:Bl ratio of approximately 5. Due to low boron retention of liposomal 

Na3[B20H19] andNa4[e 2 -B20H17OH] and rapid clearance of liposomal [2-NH3B10H9] K ,the 

enhanced retention of liposomal Na3[ae-B20H17NH3] was not due to the anionic charge or substitution 

in the borane cage. Rather, it could be attributed to their facile intracellular oxidation to an 

extremely reactive NH3-substituted [n-B20H18] 2K electrophilic anion, [B20H17NH3] K .Another 

anion [ae-B20H17NH3] 3K also was encapsulated into liposomes prepared with 5% PEG-2000distearoyl 

phosphatidylethanolamine as a constituent of the membrane. These liposomes had 

5 

7 

9 

NH 3 

SH 

H2 N 

3– 

4– 

NH 2 

+Cl – 

(CH 2 ) 3 NH 2 

3– 

BH 

B 

C 

CH 

NHCl – 

+ 

FIGURE 6.7 Structures of hydrophilic and lipophilic boron-containing compounds that have been incorporated 

into liposomes. Carboranylpropylamine (CPA, 4), [1-(2 0 -B10H9)-2-NH3B10H8] 3K (5), [n-B20H18] 2K (6), 

[B 20H 17SH] 4K (7), Na 3[a2-B 20H 17NH 2CH 2CH 2NH 2](9) and boronated water soluble acridine (WSA, 11) 

were encapsulated into the aqueous core. K[nido-7-CH3(CH2)15-7,8-C2B9H11] (8) and cholesteryl 1,12dicarba-closo-dodecaboranel-carboxylate 

(10) were incorporated into the lipid bilayer of liposomes. 



longer in vivo circulation times, which resulted in continued accumulation of boron in the tumor 

over the entire 48 h time period, and reached a maximum concentration of 47 mg B/g tumor. 

[B 20H 17SH] 4K (7), a thiol derivative of [B 20H 18] 4K , possesses a reactive thiol substituent and 

this can be oxidized into the more reactive [B 20H 17SH] 2K anion. Both of these were considered to 

be essential for high tumor boron retention 120 and they have been encapsulated into small, unilamellar 

liposomes. Biodistribution was determined after i.v. injection into BALB/c mice bearing 

EMT6 tumors. At low I.D.s, tumor boron concentrations increased throughout the duration of the 

experiment, resulting in a maximum concentration of 47 mg B/g tumor at 48 h, which corresponded 

to 22.2% I.D./g and a T:Bl ratio of 7.7. This was the most promising of the polyhedral borane anions 

that had been investigated for liposomal delivery. Although they were able to deliver adequate 

amounts of boron to tumor cells, their application to BNCT has been limited due to their low 

incorporation efficiency (approximately 3%). 

Lipophilic boron compounds incorporated into the lipid bilayer would be an alternative 

approach. Small unilamellar vesicles composed of 3:3:1 ratio of distearoylphosphatidylcholine, 

cholesterol and K[nido-7-CH3(CH2)15-7,8-C2B9H11] (8) in the lipid bilayer and Na3[a2-B20H17 

NH2CH2CH2NH2] (9) in the aqueous core were produced as a delivery agents for NCT mediated 

synovectomy. 121 Biodistribution studies were carried out in Louvain rats that had a collageninduced 

arthritis. The maximum synovial boron concentration was 29 mg/g tissue at 30 h and 

this had only decreased to 22 mg/g at 96 h following i.v. administration. The prolonged retention 

by synovium provided sufficient time for extensive clearance of boron from other tissues so that at 

96 h the synovium to blood (Syn:Bl) ratio was 3.0. To accelerate blood clearance, serum stability of 

the liposomes was lowered by increasing the proportion of K[nido-7-CH3(CH2)15-7,8-C2B9H11] 

embedded in the lipid bilayer. Liposomes were formulated with a 3:3:2 ratio of DSPC to Ch to 

K[nido-7-CH3(CH2)15-7,8-C2B9H11] in the lipid bilayer and Na3[a2-B20H17NH2CH2CH2NH2] was 

encapsulated in the aqueous core. The boron concentration in the synovium reached a maximum of 

26 mg/g at 48 h with a Syn:Bl ratio of 2, following which it slowly decreased to 14 mg/g at 96 h at 

which time the Syn:Bl ratio was 7.5. 121 

Another method to deliver hydrophilic boron containing compounds would be to incorporate 

them into cholesterol to target tumor cells expressing amplified low density lipoprotein (LDL) 

receptors. 122–124 Glioma cells, which absorb more cholesterol, have been reported to take up more 

LDL than the corresponding normal tissue cells. 125–127 The cellular uptake of liposomal cholesteryl 

1,12-dicarba-closo-dodecaboranel-carboxylate (10) by two fast growing human glioma cell lines, 

SF-763 and SF-767, was mediated via the LDL receptor and was much higher than that of human 

neurons. The cellular boron concentration was approximately 10–11 times greater than that 

required for BNCT. 128 

6.6.3 BORON DELIVERY BY TARGETED LIPOSOMES 

To improve the specificity of liposomally encapsulated drugs and to increase the amount of boron 

delivered, targeting moieties have been attached to the surface of liposomes. These could be any 

molecules that selectively recognized and bound to target antigens or receptors that were over 

expressed on neoplastic cells or tumor-associated neovasculature. These have included either intact 

mAb molecules or fragments, LMW, naturally occurring or synthetic ligands such as peptides or 

receptor-binding-ligands such as EGF. To date, liposomes linked to mAbs or their fragments, 129 

EGF, 130 folate, 131 and transferin, 112 have been the most extensively studied as targeting moieties 

(Figure 6.8). 

6.6.3.1 Immunoliposomes 

The murine anticarcinoembryonic antigen (CEA) mAb 2C-8 has been conjugated to large multilamellar 

liposomes containing 10 B compounds. 132,133 The maximum number of 10 B atoms attached 


90 

A 

per molecule of mAb was approximately 1.2!10 4 . These immunoliposomes bound selectively to 

the human pancreatic carcinoma cell line, AsPC-1 that overexpressed CEA. Incubating the immunoliposomes 

with either MRKnu/nu-1 or AsPC-1 tumor cells, suppressed in vitro tumor cell growth 

following thermal neutron irradiation. 134 This was dependent upon the liposomal concentration of 

the 10 B-compound and on the number of molecules of mAb conjugated to the liposomes. Immunoliposomes 

containing either (Et4N)2B10H10 and linked to the mAb MGb 2, directed against 

human gastric cancer 135,136 or water-soluble boronated acridine (WSA, 11) linked to trastuzumab, 

directed against HER-2, have been prepared and evaluated in vitro. 129 There was specific binding 

and high uptake of these immunoliposomes, which delivered a sufficient amount of 10 B to produce a 

tumoricidal effect following thermal neutron irradiation. 

6.6.3.2 Folate Receptor-Targeted Liposomes 

C 

H + 

B 

A highly ionized boron compound, Na3B20H17NH3, was incorporated into liposomes by passive 

loading. 131,137,138 This showed high in vitro uptake by the FR expressing human cell line KB 

(American Type Culture Collection CCL 17), which originally was thought to be derived from a 

squamous cell carcinoma of the mouth, and subsequently was shown to be identical to HeLa cells, 

as determined by isoenyzyme markers, DNA fingerprinting, and karyotypic analysis. KB tumorbearing 

mice that received either FR-targeted or nontargeted control liposomes had equivalent 

tumor boron values (approximately 85 mg/g), which attained a maximum at 24 h, while the T:Bl 

ratio reached a maximum at 72 h. Additional studies were carried out with the lipophilic boron 

compound, K[nido-7-CH3(CH2)15-7,8-C2B9H11. This was incorporated into large unilamellar 

vesicles, approximately 200 nm in diameter, which were composed of egg PC/Chol/K[nido-7- 

CH3(CH2)15-7,8-C2B9H11] at a 2:2:1, mol/mol ratio, and an additional 0.5 mol% of folate– 

PEG–DSPE or PEG–DSPE for the FR-targeted or nontargeted liposomal formulations. 139 The 

boron uptake by FR-overexpressing KB cells, treated with these targeted liposomes, was approximately 

10 times greater compared with those treated with control liposomes. In addition, BSH and 

D 

H + H + 


Folate 

PEG 

Liposome 

Folate receptor 

Boron compound 

FIGURE 6.8 Proposed mechanism of intracellular boron delivery based on receptor-mediated tumor cell 

targeting. Liposomes loaded with a boron compound (A) are conjugated to a targeting ligand (e.g., folate, 

transferrin (TF), or anti-EGFR antibody). These bind to receptors on the cell surface (B), following which they 

are internalized by receptor mediated endocytosis (C) into the acidified endosomal/lysosomal compartment. 

The boron compound then is released (D) into the cytosol by liposomal degredation, endosome/lysosome 

disruption or liposome–endosome/lysome membrane fusion. 



H 2 N 

H 2 N 

12 

N 

H 2 N N N H 

14 15 

N N 

16 

NH 2 · 3HCl 

NH 2 · 4HCl 

five weakly basic boronated polyamines were evaluated (Figure 6.9). Two of these were the 

spermidine derivatives N 5 -(4-carboranylbutyl)spermidine$3HCl (12) and N 5 -[4-(2-aminoethyl-ocarboranyl)butyl] 

spermidine$4HCl (13). Three were the spermine derivatives N 5 -(4-o-carboranylbutyl) 

spermine$4HCl (14), N 5 -[4-(2-aminoethyl-o-carboranyl) butyl] spermine$5HCl (15), and 

N 5 ,N 10 -bis(4-o-carboranylbutyl) spermine$4HCl (16). These were incorporated into liposomes 

by a pH-gradient-driven remote loading method with varying loading efficiencies that were influenced 

by the specific trapping agent and the structure of the boron compound. Greater loading 

efficiencies were obtained with lower molecular-weight boron derivatives, using ammonium sulfate 

as the trapping agent, compared to those obtained with sodium citrate. 

6.6.3.3 EGFR-Targeted Liposomes 

NH 2 · 4HCl 

Acridine is a WS, DNA-intercalator. Its boronated derivative WSA was incorporated into liposomes 

composed of EGF-conjugated lipids. Their surface contained approximately 5 mol% PEG 

and 10–15 molecules of EGF, and 10 4 –10 5 of WSA molecules were encapsulated. These liposomes 

had EGFR-specific cellular binding to cultured human glioma cells 130,140 and were internalized 

following specific binding to the receptor. Following internalization, WSA primarily was localized 

H 2 N 

H 2 N 

H 2 N · HCl 

H 2 N · HCl 

N 

13 

N 

BH 

C 

CH 

H 

N 

NH 2 · 3HCl 

NH 2 · 4HCl 

FIGURE 6.9 Structures of five weakly basic boronated polyamines encapsulated in FR-targeting liposomes. 

Two spermidine (12, 13) and three spermine derivatives (14–16) that contain hydrophilic amine groups and 

lipophilic carboranyl cages had DNA-binding properties. 


92 

in the cytoplasm, and had high cellular retention with 80% of the boron remaining cell-associated 

after 48 h. 141 

6.7 BORON DELIVERY BY DEXTRANS 

Dextrans are glucose polymers that consist mainly of a linear a-1,6-glucosidic linkage with some 

degree of branching via a 1,3-linkage. 142,143 Dextrans have been used extensively as drug and 

protein carriers to increase drug circulation time. 144,145 In addition, native or chemically-modified 

dextrans have been used for passive targeting to tumors, the RES or active receptor-specific cellular 

targeting. To link boron compounds to dextrans, 146 b-decachloro-o-carborane derivatives, in which 

one of the carbon atoms was substituted by –CH2CHOHCH2–O–CH2CHaCH2, were epoxidized 

and then subsequently bound to dextran with a resulting boron content of 4.3% (w/w). 147 The 

modified dextran then could be attached to tumor-specific antibodies. 147–150 BSH was covalently 

coupled to dextran derivatives by two methods. 151 In the first method, dextran was activated with 

1-cyano-4-(dimethylamino)pyridine (CDAP) and subsequently coupled with 2-aminoethyl pyridyl 

disulfide. Then, thiolated dextran was linked to BSH in a disulfide exchange reaction. A total of 10– 

20 boron cages were attached to each dextran chain. In the second method, dextran was derivatized 

to a multiallyl derivative (Figure 6.10, 17), which was reacted with BSH in a free-radical-initiated 

addition reaction. Using this method, 100–125 boron cages could be attached per dextran chain, 

suggesting that this derivative might be a promising template for the development of other HMW 

delivery agents. In the second method, designed to target EGFR overexpressing cells, EGF and 

BSH were covalently linked to a 70 kDa dextran (18). 152–154 Bioconjugates, having a small number 

of BSH molecules, attained maximum in vitro binding at 4 h with the human glioma cell line U-343 

O 

O 

O 

OH 

O 

OH 

O O 

OH 

O O 

OH O 

OH 

OH O 

OH O 

OH O 

OH O 

OH 

O O 

OH 

BSH or EGF-SH, 2 or 5h, 50 o C 

(NH 4 ) 2 S 2 O 8 /K 2 S 2 O 8 

O 

O 

O 

OH 

O 

OH 

O O 

OH 

O O 

S 

OH O S 

OH 

EGF 

OH O 

OH O 

17 18 

OH O 

OH O 

OH 

O O 

OH 

FIGURE 6.10 Preparation of EGF-targeted, boronated dextrans. The bioconjugate was prepared by a freeradical-initiated 

addition reaction between multiallyl dextran derivatives and BSH or thiolated EGF at 508C 

using K2S2O8 as an initiator. 



S 

BH 

B


MGaC12:6. In contrast, there was a slow increase of binding over 24 h for those having a large 

number of BSH molecules. Although most of the bioconjugates were internalized, in vitro retention 

was low, as was in vivo uptake following i.v. injection into nude mice bearing s.c. implants of 

Chinese hamster ovary (CHO) cells transfected with the human gene encoding EGFR (designated 

CHO–EGFR). However, following i.t. injection, boron uptake was higher with CHO–EGFR(C) 

tumors compared to wildtype EGFR(K) CHO tumors. 155 

6.8 OTHER MACROMOLECULES USED FOR DELIVERING 

BORON COMPOUNDS 

Polylysine is another polymer having multiple reactive amino groups that has been used as a 

platform for the delivery of boron compounds. 53,156 The protein-binding polyhedral boron derivatives, 

isocyanatoundecahydro-closo-dodecaborate (B12H11NCO 2K ), was linked to polylysine and 

subsequently to the anti-B16 melanoma mAb IB16-6 using two heterobifunctional linkers, SPDP 

and sulfo-MBS. The bioconjugate had an average of 2700 boron atoms per molecule and retained 

58% of the immunoreactivity of the native antibody, as determined by a semiquantitative immunofluorescent 

assay or by ELISA. Other bioconjugates prepared by this method had greater than 1000 

boron atoms per molecule of antibody and retained 40–90% of the immunoreactivity of the native 

antibody. 53 Using another approach, site-specific linkage of boronated polylysine to the carbohydrate 

moieties of anti-TSH antibody resulted in a bioconjugate that had approximately 6!10 3 

boron atoms with retention of its immunoreactivity. 156 

A streptavidin/biotin system also has been developed to specifically deliver boron to tumors. 

Biotin was linked to a mAb and streptavidin was attached to the boron containing moiety. The 

indirect linking of boron to the mAb minimized loss of its immunoreactivity. BSH was attached to 

poly-(D-glutamate D-lysine) (poly-GL) via a heterobifunctional agent. 157 This boronated poly-GL 

then was activated by a carbodiimide reagent and in turn reacted with streptavidin. Another 

approach employed a streptavidin mutant that had 20 cysteine residues per molecule. BSH was 

conjugated via sulfhydryl-specific bifunctional reagents to incorporate approximately 230 boron 

atoms/molecule. 158 A closomer species with an icosahedral dodecaborate core and twelve pendant 

anionic nido-7,8-carborane groups was developed as a new class of unimolecular nanovehicles for 

evaluation as a delivery agent for BNCT. 159 

6.9 DELIVERY OF BORON-CONTAINING MACROMOLECULES 

TO BRAIN TUMORS 

6.9.1 GENERAL CONSIDERATIONS 

Drug delivery to brain tumors is dependent upon (1) the plasma concentration profile of the drug, 

which depends upon the amount and route of administration; (2) the ability of the agent to traverse 

the BBB; (3) blood flow within the tumor; and (4) the lipophilicity of the drug. In general, a high 

steady-state blood concentration will maximize brain uptake, while rapid clearance will reduce it, 

except in the case of intra-arterial (i.a.) drug administration. Although the i.v. route currently is 

being used clinically to administer both BSH and BPA, this may not be ideal for boron containing 

macromolecules and other strategies must be employed to improve their delivery. 

6.9.2 DRUG-TRANSPORT VECTORS 

One approach to improve brain tumor uptake of boron compounds has been to conjugate them to a 

drug-transport vector by means of receptor-specific transport systems. 160,161 Proteins such as 

insulin, insulin-like growth factor (IGF), TF, 162 and leptin can traverse the BBB. BSH encapsulated 


94 

in TF–PEG liposomes had a prolonged residence time in the circulation and low RES uptake in 

tumor-bearing mice, resulting in enhanced extravasation of the liposomes into the tumor and 

concomitant internalization by receptor-mediated endocytosis. 163,164 Mice that received BSH 

containing TF-liposomes following by BNCT had a significant prolongation in survival time 

compared to those that received PEG-liposomes, bare liposomes and free BSH, thereby establishing 

proof-of-principle for transcytosis of a boron containing nanovehicle. 112 

6.9.3 DIRECT INTRACEREBRAL DELIVERY 

Studies carried out by us have clearly demonstrated that the systemic route of administration is not 

suitable for delivery of boronated EGF or mAbs to glioma-bearing rats. 75,165 Intravenous injection 

of technetium-99m labeled EGF to rats bearing intracerebral implants of the C6 rat glioma, which 

had been genetically engineered to express the human EGFR gene, resulted in 0.14% I.D. localizing 

in the tumor. Intracarotid (i.c.) injection with or without BBB disruption increased the tumor uptake 

from 0.34 to 0.45% I.D./g, but based even on the most optimistic assumptions the amount of boron 

that could be delivered to the tumor by i.v. injection, this would have been inadequate for BNCT. 165 

Direct i.t. injection of boronated EGF (BSD–EGF), on the other hand, resulted in tumor boron 

concentrations of 22 mg/g compared to 0.01 mg/g following i.v. injection and almost identical boron 

uptake values were obtained using the F98 EGFR glioma model. 77 This was produced by transfecting 

F98 glioma cells with the gene encoding human EGFR. Based on our biodistribution results, 

therapy studies were initiated with the F98EGFR glioma in syngeneic Fischer rats. F98EGFR 

glioma bearing rats that received BSD–EGF i.t. had a MST of 45Gd compared to 33G2 din 

animals that had EGFR(K) wildtype F98 gliomas. Because it is unlikely that any single boron 

delivery agent will be able to target all tumor cells, the combination of i.t. administration of 

BSD–EGF with i.v. injection of BPA was evaluated. This resulted in further increase in MST to 

57G8 d compared to 39G2 d for i.v. BPA alone. 73 These data provide proof-of-principle for the 

idea of using a combination of LMW and HMW boron delivery agents. 

6.9.4 CONVECTION-ENHANCED DELIVERY 


Convection-Enhanced Delivery (CED), by which therapeutic agents are directly infused into the 

brain, is an innovative method to increase their uptake and distribution. 166–168 Under normal 

physiological conditions, interstitial fluids move through the brain by both convection and diffusion. 

Diffusion of a drug in tissue depends upon its molecular weight, ionic charge and its 

concentration gradient within normal tissue and the tumor. The higher the molecular weight of 

the drug, the more positively charged the ionic species and the lower its concentration, the slower 

its diffusion. For example, diffusion of antibody into a tumor requires 3 d to diffuse 1 mm from the 

point of origin. Unlike diffusion, however, convection or “bulk” flow results from a pressure 

gradient that is independent of themolecularweightofthesubstance. CED potentially can 

improve the targeting of both LMW and HMW molecules, as well as liposomes, to the central 

nervous system by applying a pressure gradient to establish bulk flow during interstitial infusion. 

The volume of distribution (V d) is a linear function of the volume of the infusate (V i). CED has been 

used to efficiently deliver drugs and HMW agents such as mAbs and toxin fusion proteins to brain 

tumors. 168–170 CED can provide more homogenous dispersion of the agent and at higher concentrations 

than otherwise would be attainable by i.v. injection. 165 For example, in our own studies, 

CED of 125 I-labeled EGF to F98EGFR glioma-bearing rats resulted in 47% I.D./g of the bioconjugate 

localizing in the tumor compared to 10% I.D./g in normal brain at 24 h following administration. 

The corresponding boron values were 22 and 2.9–4.9 mg/g, respectively. 76 Based on these results, 

therapy studies were initiated. F98EGFR glioma-bearing rats that received BD–EGF by CED had a 

MST of 53G13 d compared to 40G5 d for animals that received BPA i.v. 73 Similar studies have 

been carried out using either boronated cetuximab (IMC-C225) or the mAb L8A4, 171,172 which is 



specifically directed against the tumor-specific mutant isoform, EGFRvIII, and comparable results 

were obtained. 173 Direct intracerebral administration of these and other HMW agents by CED has 

opened up the possibility that they actually could be used clinically, since CED is being used to 

administer radiolabeled antibodies, toxin fusion proteins, and gene vectors to patients with GBM. It 

is only a matter of time before this approach also will be used to deliver both LMW and HMW 

boron-containing agents for NCT. 

6.10 CLINICAL CONSIDERATIONS AND CONCLUSIONS 

In this review we have focused on HMW boron delivery agents and nanovehicles that potentially 

could be used clinically for targeting intra- and extracranial tumors. Studies carried out in glioma 

bearing rats have demonstrated that boronated EGF and the mAb, cetuximab, both of which bind to 

EGFR, selectively targeted receptor (C) tumors following direct i.c. delivery. Furthermore, 

following BNCT, a significant increase in MST was observed, and this was further enhanced if 

BPA was administered in combination with the HMW agents. These studies provide proof-of-principle 

first for the potential utility of HMW agents, and second, the therapeutic gain associated with 

the combination of HMW and LMW boron delivery agents. 

There is a question as to whether or not any of these agents will ever be used clinically. Critical 

issues must be addressed if BNCT is to ever become a useful modality for the treatment of cancer. 

Large clinical trials, preferably randomized, must be carried out to convincingly demonstrate the 

efficacy of BPA and BSH, the two drugs that currently are being used. Once efficacy has been 

established, studies with HMW EGFR targeting agents could move forward. 

Both direct i.t. injection 174,175 and CED 170,176–179 have been used clinically to deliver mAbs 

and toxin fusion proteins to patients who have had surgical resection of their brain tumors. These 

studies provide a strong clinical rationale for the direct intracerabral delivery of HMW agents. 

Initially, the primary focus should be on determining the safety of administering them to patients 

prior to surgical resection of their brain tumors. After this has been established, then biodistribution 

studies could be carried out in patients scheduled to undergo surgical resection of their brain 

tumors. The patients’ tumor and normal tissues would then be analyzed for their boron content, 

and if there was evidence of preferential tumor localization with boron concentrations in the range 

10–20 mg/g and normal brain concentrations of less than 5 mg/g, then therapy studies could 

be undertaken. 

Because there is considerable variability in EGFR expression in gliomas, it is highly unlikely 

that any single agent will be able to deliver the requisite amount of boron to all tumor cells, and that 

HMW agents would need to be used in combination with BPA/BSH. This general plan would also 

be applicable to the other HMW delivery agents and nanovehicles that have been discussed in this 

review. The joining together of chemistry and nanotechnology 180,181 represents a major step 

forward for the development of effective boron delivery agents for NCT. Nanovehicles offer the 

possibility of tumor targeting with enhanced boron payloads. Potentially, this could solve the 

central problem of how to selectively deliver large number of boron atoms to individual 

cancer cells. 

As can be seen from the preceding discussion, the development of HMW boron delivery agents 

must proceed in step with strategies to optimize their delivery and an appreciation as to how they 

would be used clinically. Intracerebral delivery has been used in clinically advanced settings, but 

nuclear reactors, which currently are the only source of neutrons for BNCT, would not be conducive 

to this. Therefore, the development of accelerator-neutron sources, 182 which could be easily 

sited in hospitals, is especially important. This also would facilitate the initiation of large scale 

clinical trials at selected centers that treat large numbers of patients with brain tumors and would 

permit evaluation of new boron delivery agents. 


96 

In conclusion, there is a plethora of HMW boron delivery agents that have been designed and 

synthesized. The challenge is to move from experimental animal studies to clinical biodistribution 

studies, a step which has yet to be taken. 

6.11 SUMMARY 

BNCT is based on the nuclear capture and fission reactions that occur when nonradioactive boron- 

10 is irradiated with low energy thermal neutrons to yield LET alpha particles ( 4 He) and recoiling 

lithium-7 ( 7 Li) nuclei. For BNCT to be successful, a sufficient number of 10 B atoms (approximately 

10 9 atoms/cell) must be selectively delivered to the tumor and enough thermal neutrons must be 

absorbed by them to sustain a lethal 10 B(n,a) 7 Li capture reaction. BNCT primarily has been used to 

treat patients with brain tumors, and more recently those with head-and-neck cancer. Two LMW 

boron delivery agents currently are being used clinically, BSH and BPA. However, a variety of 

HMW agents consisting of macromolecules and nanovehicles have been developed. This review 

focuses on the latter, which includes mAbs, dendrimers, liposomes, dextrans, polylysine, avidin, 

FA, and both epidermal and vascular endothelial growth factors (EGF and VEGF). Procedures for 

introducing boron atoms into these HMW agents and their chemical properties are discussed. 

In vivo studies on their biodistribution are described, and the efficacy of a subset of them, those 

which have been used for BNCT of tumors in experimental animals, will be discussed. 

Because brain tumors currently are the primary candidates for treatment by BNCT, delivery of 

these HMW agents across the BBB presents a special challenge. Various routes of administration are 

discussed, including receptor-facilitated transcytosis following i.v. administration, direct i.t injection 

and convection-enhanced delivery in which a pump is used to apply a pressure gradient to establish 

bulk flow of the HMW agent during interstitial infusion. Finally, we have concluded with a discussion 

relating to issues that must be addressed if these HMW agents are to be used clinically. 

ACKNOWLEDGMENTS 

We thank Dr. Achintya Bandyopadhyaya for suggestions on figures and Mrs. Beth Kahl for 

secretarial assistance. We thank Bentham Science Publishers, Ltd. for granting copyright release 

to republish the text and figures in this chapter. 

REFERENCES 


1. Berger, M. S., Malignant astrocytomas: Surgical aspects, Semin. Oncol., 21, 172, 1994. 

2. Gutin, P. H. and Posner, J. B., Neuro-oncology: Diagnosis and management of cerebral gliomas— 

past, present, and future, Neurosurgery, 47, 1, 2000. 

3. Parney, I. F. and Chang, S. M., Current chemotherapy for glioblastoma, Cancer J., 9, 149, 2003. 

4. Paul, D. B. and Kruse, C. A., Immunologic approaches to therapy for brain tumors, Curr. Neurol. 

Neurosci. Rep., 1, 238, 2001. 

5. Rainov, N. G. and Ren, H., Gene therapy for human malignant brain tumors, Cancer J., 9, 180, 2003. 

6. Curran, W.J., Jr., et al., Recursive partitioning analysis of prognostic factors in three Radiation 

Therapy Oncology Group malignant glioma trials, J. Natl Cancer Inst., 85, 704, 1993. 

7. Lacroix, M. et al., A multivariate analysis of 416 patients with glioblastoma multiforme: Prognosis, 

extent of resection, and survival, J. Neurosurg., 95, 190, 2001. 

8. Hentschel, S. J. and Lang, F. F., Current surgical management of glioblastoma, Cancer J., 9, 113, 

2003. 

9. Laws, E.R., Jr. and Shaffrey, M. E., The inherent invasiveness of cerebral gliomas: Implications for 

clinical management, Int. J. Dev. NeuroSci., 17, 413, 1999. 

10. Ware, M. L., Berger, M. S., and Binder, D. K., Molecular biology of glioma tumorigenesis, Histol. 

Histopathol., 18, 207, 2003. 



11. Parney, I. F., Hao, C., and Petruk, K. C., Glioma immunology and immunotherapy, Neurosurgery, 

46, 778, 2000. 

12. Kaczarek, E. et al., Dissecting glioma invasion: Interrelation of adhesion, migration and intercellular 

contacts determine the invasive phenotype, Int. J. Dev. NeuroSci., 17, 625, 1999. 

13. Nutt, C. L., Matthews, R. T., and Hockfield, S., Glial tumor invasion: A role for the upregulation and 

cleavage of BEHAB/brevican, Neuroscientist, 7, 113, 2001. 

14. Halperin, E. C., Burger, P. C., and Bullard, D. E., The fallacy of the localized supratentorial 

malignant glioma, Int. J. Radiat. Oncol. Biol. Phys., 15, 505, 1988. 

15. Barth, R. F., A critical assessment of boron neutron capture therapy: An overview, J. Neurooncol., 

62, 1, 2003. 

16. Nakagawa, Y. et al., Clinical review of the Japanese experience with boron neutron capture therapy 

and a proposed strategy using epithermal neutron beams, J. Neurooncol., 62, 87, 2003. 

17. Wadabayashi, N. et al., Selective boron accumulation in human ocular melanoma vs. surrounding 

eye components after 10 B1-p-boronophenylalanine administration. Prerequisite for clinical trial of 

neutron-capture therapy, Melanoma Res., 4, 185, 1994. 

18. Busse, P. M. et al., A critical examination of the results from the Harvard-MIT NCT program phase I 

clinical trial of neutron capture therapy for intracranial disease, J. Neurooncol., 62, 111, 2003. 

19. Kato, I. et al., Effectiveness of BNCT for recurrent head and neck malignancies, Appl. Radiat. Isot., 

61, 1069, 2004. 

20. Rao, M. et al., BNCT of 3 cases of spontaneous head and neck cancer in feline patients, Appl. Radiat. 

Isot., 61, 947, 2004. 

21. Koivunoro, H. et al., BNCT dose distribution in liver with epithermal D–D and D–T fusion-based 

neutron beams, Appl. Radiat. Isot., 61, 853, 2004. 

22. Pinelli, T. et al., TAOrMINA: from the first idea to the application to the human liver, In Research 

and Development in Neutron Capture Therapy, Sauerwein, M. W., Moss, R., Wittig, A. et al., Eds., 

Modduzzi Editore, International Proceedings Division, Bologna, pp. 1065–1072, 2002. 

23. Coderre, J. A. et al., Boron neutron capture therapy: Cellular targeting of high linear energy transfer 

radiation, Technol. Cancer Res. Treat., 2, 355, 2003. 

24. Barth, R. F. et al., Boron neutron capture therapy of cancer: Current status and future prospects, Clin. 

Cancer Res., 11, 3987, 2005. 

25. Zamenhof, R. G. et al., Eleventh World Congress on Neutron Capture Therapy, Appl. Radiat. Isot., 

61, 731, 2004. 

26. Farr, L. E. et al., Neutron capture therapy with boron in the treatment of glioblastoma multiforme, 

Am. J. Roentgenol. Radium Ther. Nucl. Med., 71, 279, 1954. 

27. Goodwin, J. T. et al., Pathological study of eight patients with glioblastoma multiforme treated by 

neutron-capture therapy using boron 10, Cancer, 8, 601, 1955. 

28. Snyder, H. R., Reedy, A. J., and Lennarz, W. J., Synthesis of aromatic boronic acids. Aldehydo 

boronic acids and a boronic acid analog of tyrosine, J. Am. Chem. Soc., 80, 835, 1958. 

29. Soloway, A. H., Hatanaka, H., and Davis, M. A., Penetration of brain and brain tumor. VII. Tumorbinding 

sulfhydryl boron compounds, J. Med. Chem., 10, 714, 1967. 

30. Hatanaka, H. and Nakagawa, Y., Clinical results of long-surviving brain tumor patients who underwent 

boron neutron capture therapy, Int. J. Radiat. Oncol. Biol. Phys., 28, 1061, 1994. 

31. Mishima, Y., Selective thermal neutron capture therapy of cancer cells using their specific metabolic 

activities—melanoma as prototype, In Cancer Neutron Capture Therapy, Mishima, Y., Ed., Plenum 

Press, New York, pp. 1–26, 1996. 

32. Ang, K. K. et al., Impact of epidermal growth factor receptor expression on survival and pattern of 

relapse in patients with advanced head and neck carcinoma, Cancer Res., 62, 7350, 2002. 

33. Hawthorne, M. F. and Lee, M. W., A critical assessment of boron target compounds for boron 

neutron capture therapy, J. Neurooncol., 62, 33, 2003. 

34. Soloway, A. H. et al., The chemistry of neutron capture therapy, Chem. Rev., 98, 1515, 1998. 

35. Gillies, E. R. and Frechet, J. M. J., Dendrimers and dendritic polymers in drug delivery, Drug 

Discov. Today, 10, 35, 2005. 

36. Esfand, R. and Tomalia, D. A., Poly(amidoamine) (PAMAM) dendrimers: From biomimicry to drug 

delivery and biomedical applications, Drug Discov. Today, 6, 427, 2001. 


98 


37. Tomalia, D.A., Birth of a new macromolecular architecture: Dendrimers as quantized building 

blocks for nanoscale synthetic organic chemistry, Aldrichimica ACTA, 37, 39, 2004 

38. Verheyde, B., Maes, W., and Dehaen, W., The use of 1,3,5-triazines in dendrimer synthesis, Mater. 

Sci. Eng., C, 18, 243, 2001. 

39. McCarthy, T. D. et al., Dendrimers as drugs: Discovery and preclinical and clinical development of 

dendrimer-based microbicides for HIV and STI prevention, Mol. Pharmacol., 2, 312, 2005. 

40. Venditto, V. J., Regino, C. A., and Brechbiel, M. W., PAMAM dendrimer based macromolecules as 

improved contrast agents, Mol. Pharmacol., 2, 302, 2005. 


release and targeted delivery, Mol. Pharmacol., 2, 264, 2005. 

42. Majoros, I. J. et al., Poly(amidoamine) dendrimer-based multifunctional engineered nanodevice for 

cancer therapy, J. Med. Chem., 48, 5892, 2005. 

43. Boas, U. and Heegaard, P. M., Dendrimers in drug research, Chem. Soc. Rev., 33, 43, 2004. 

44. Klajnert, B. and Bryszewska, M., Dendrimers: Properties and applications, Acta Biochim. Pol., 48, 

199, 2001. 

45. Sharkey, R. M. and Goldenberg, D. M., Perspectives on cancer therapy with radiolabeled monoclonal 

antibodies, J. Nucl. Med., 46(Suppl 1), 115S, 2005. 

46. Jaracz, S. et al., Recent advances in tumor-targeting anticancer drug conjugates, Bioorg. Med. 

Chem., 13, 5043, 2005. 

47. Garnett, M. C., Targeted drug conjugates: Principles and progress, Adv. Drug Deliv. Rev., 53, 171, 

2001. 

48. Chari, R. V., Targeted delivery of chemotherapeutics: Tumor-activated prodrug therapy, Adv. Drug 

Deliv. Rev., 31, 89, 1998. 

49. Hamblett, K. J. et al., Effects of drug loading on the antitumor activity of a monoclonal antibody drug 

conjugate, Clin. Cancer Res., 10, 7063, 2004. 

50. Trail, P. A., King, H. D., and Dubowchik, G. M., Monoclonal antibody drug immunoconjugates for 

targeted treatment of cancer, Cancer Immunol. Immunother., 52, 328, 2003. 

51. Fracasso, G. et al., Immunotoxins and other conjugates: Preparation and general characteristics, 

Mini-Rev. Med. Chem., 4, 545, 2004. 

52. Liu, L. et al., Bispecific antibodies as targeting agents for boron neutron capture therapy of brain 

tumors, J. Hematother., 4, 477, 1995. 

53. Alam, F. et al., Boron neutron capture therapy: Linkage of a boronated macromolecule to monoclonal 

antibodies directed against tumor-associated antigens, J. Med. Chem., 32, 2326, 1989. 

54. Barth, R. F. et al., Neutron capture using boronated monoclonal antibody directed against tumorassociated 

antigens, Cancer Detect. Prev., 5, 315, 1982. 

55. Alam, F., Barth, R. F., and Soloway, A. H., Boron containing immunoconjugates for neutron capture 

therapy of cancer and for immunocyto chemistry, Antibody Immunoconjugates Radiopharm., 2, 145, 

1989. 

56. Tolpin, E. I. et al., Boron neutron capture therapy of cerebral gliomas. II. Utilization of the blood– 

brain barrier and tumor-specific antigens for the selective concentration of boron in gliomas, 

Oncology, 32, 223, 1975. 

57. Sneath, R. L. et al., Protein-binding polyhedral boranes, J. Med. Chem., 19, 1290, 1976. 

58. Varadarajan, A. et al., Conjugation of phenyl isothiocyanate derivatives of carborane to antitumor 

antibody and in vivo localization of conjugates in nude mice, Bioconjug. Chem., 2, 102, 1991. 

59. Takahashi, T. et al., Preliminary study for application of anti-alpha-fetoprotein monoclonal antibody 

to boron-neutron capture therapy, Jpn. J. Exp. Med., 57, 83, 1987. 

60. Compostella, F. et al., Synthesis of glycosyl carboranes with different linkers between the sugar and 

the boron cage moieties, In Research and Development in Neutron Capture Therapy. Proceedings of 

the 10th International Congress on Neutron Capture Therapy, Saverwein, W., Moss, R., and Wittig, 

A., Eds., Monduzzi Editore, Bologna, p.8, 2002. 

61. Giovenzana, G. B. et al., Synthesis of carboranyl derivatives of alkynyl glycosides as potential 

BNCT agents, Tetrahedron, 55, 14123, 1999. 

62. Alam, F. et al., Dicesium N-succinimidyl 3-(undecahydro-closo-dodecaboranyldithio)propionate, a 

novel heterobifunctional boronating agent, J. Med. Chem., 28, 522, 1985. 



63. Barth, R. F. et al., Boronated starburst dendrimer-monoclonal antibody immunoconjugates: evaluation 

as a potential delivery system for neutron capture therapy, Bioconjug. Chem., 5, 58, 1994. 

64. Liu, L. et al., Critical evaluation of bispecific antibodies as targeting agents for boron neutron capture 

therapy of brain tumors, Anticancer Res., 16, 2581, 1996. 

65. Barth, R. F. et al., In vivo distribution of boronated monoclonal antibodies and starburst dendrimers, 

In Adv Neutron Capture Ther, Soloway, A. H., Barth, R. F., and Carpenter, D.E., Eds., Plenum Press, 

New York, p. 35, 1993. 

66. Wu, G. et al., Site-specific conjugation of boron-containing dendrimers to anti-EGF receptor monoclonal 

antibody cetuximab (IMC-C225) and its evaluation as a potential delivery agent for neutron 

capture therapy, Bioconjug. Chem., 15, 185, 2004. 

67. Barth, R. F. et al., Neutron capture therapy of epidermal growth factor (C) gliomas using boronated 

cetuximab (IMC-C225) as a delivery agent, Appl. Radiat. Isot., 61, 899, 2004. 

68. Arteaga, C. L., Overview of epidermal growth factor receptor biology and its role as a therapeutic 

target in human neoplasia, Semin. Oncol., 29, 3, 2002. 

69. Mendelson, J. and Beselga, J., Epidermal growth factor receptor targeting in cancer, Seminars 

Oncol., 33, 369, 2006. 

70. Pal, S. K. and Pegram, M., Epidermal growth factor receptor and signal transduction: potential 

targets for anti-cancer therapy, Anticancer Drugs, 16, 483, 2005. 

71. Normanno, N. et al., The ErbB receptors and their ligands in cancer: an overview, Curr. Drug 

Targets, 6, 243, 2005. 

72. Jorissen, R. N. et al., Epidermal growth factor receptor: mechanisms of activation and signaling, Exp. 

Cell Res., 284, 31, 2003. 

73. Yang, W. et al., Boronated epidermal growth factor as a delivery agent for neutron capture therapy of 

EGF receptor positive gliomas, Appl. Radiat. Isot., 61, 981, 2004. 

74. Capala, J. et al., Boronated epidermal growth factor as a potential targeting agent for boron neutron 

capture therapy of brain tumors, Bioconjug. Chem., 7, 7, 1996. 

75. Yang, W. et al., Intratumoral delivery of boronated epidermal growth factor for neutron capture 

therapy of brain tumors, Cancer Res., 57, 4333, 1997. 

76. Yang, W. et al., Convection-enhanced delivery of boronated epidermal growth factor for molecular 

targeting of EGF receptor-positive gliomas, Cancer Res., 62, 6552, 2002. 

77. Barth, R. F. et al., Molecular targeting of the epidermal growth factor receptor for neutron capture 

therapy of gliomas, Cancer Res., 62, 3159, 2002. 

78. Reddy, J. A., Allagadda, V. M., and Leamon, C. P., Targeting therapeutic and imaging agents to 

folate receptor positive tumors, Curr. Pharm. Biotechnol., 6, 131, 2005. 

79. Leamon, C. P. and Reddy, J. A., Folate-targeted chemotherapy, Adv. Drug Deliv. Rev., 56, 1127, 

2004. 

80. Sudimack, J. and Lee, R. J., Targeted drug delivery via the folate receptor, Adv. Drug Deliv. Rev., 41, 

147, 2000. 

81. Paulos, C. M. et al., Ligand binding and kinetics of folate receptor recycling in vivo: impact on 

receptor-mediated drug delivery, Mol. Pharmacol., 66, 1406, 2004. 

82. Reddy, J. A. and Low, P. S., Folate-mediated targeting of therapeutic and imaging agents to cancers, 

Crit. Rev. Ther. Drug Carrier Syst., 15, 587, 1998. 

83. Stephenson Stacy, M. et al., Folate receptor-targeted liposomes as possible delivery vehicles for 

boron neutron capture therapy, Anticancer Res., 23, 3341, 2003. 

84. Ward, C. M., Folate-targeted non-viral DNA vectors for cancer gene therapy, Curr. Opin. Mol. Ther., 

2, 182, 2000. 

85. Gottschalk, S. et al., Folate receptor mediated DNA delivery into tumor cells: Protosomal disruption 

results in enhanced gene expression, Gene Ther., 1, 185, 1994. 

86. Hilgenbrink, A. R. and Low, P. S., Folate receptor-mediated drug targeting: from therapeutics to 

diagnostics, J. Pharm. Sci., 94, 2135, 2005. 

87. Shukla, S. et al., Synthesis and biological evaluation of folate receptor-targeted boronated PAMAM 

dendrimers as potential agents for neutron capture therapy, Bioconjug. Chem., 14, 158, 2003. 

88. Benouchan, M. and Colombo, B. M., Anti-angiogenic strategies for cancer therapy (Review), Int. 

J. Oncol., 27, 563, 2005. 


100 


89. Tortora, G., Melisi, D., and Ciardiello, F., Angiogenesis: a target for cancer therapy, Curr. Pharm. 

Des., 10, 11, 2004. 

90. Brekken, R. A., Li, C., and Kumar, S., Strategies for vascular targeting in tumors, Int. J. Cancer, 100, 

123, 2002. 

91. Gaya, A. M. and Rustin, G. J., Vascular disrupting agents: a new class of drug in cancer therapy, 

Clin. Oncol. (R. Coll. Radiol.), 17, 277, 2005. 

92. Bergsland, E. K., Vascular endothelial growth factor as a therapeutic target in cancer, Am. J. Health 

Syst. Pharm., 61, S4, 2004. 

93. Hicklin, D. J. and Ellis, L. M., Role of the vascular endothelial growth factor pathway in tumor 

growth and angiogenesis, J. Clin. Oncol., 23, 1011, 2005. 

94. Backer, M. V. et al., Vascular endothelial growth factor selectively targets boronated dendrimers to 

tumor vasculature, Mol. Cancer Ther., 4, 1423, 2005. 

95. Barth, R. F., Yang, W., and Coderre, J. A., Rat brain tumor models to assess the efficacy of boron 

neutron capture therapy: a critical evaluation, J. Neurooncol., 62, 61, 2003. 

96. Parrott, M. C. et al., Synthesis and properties of carborane-functionalized aliphatic polyester dendrimers, 

J. Am. Chem. Soc., 127, 12081, 2005. 

97. Park, J. W., Benz, C. C., and Martin, F. J., Future directions of liposome- and immunoliposomebased 

cancer therapeutics, Semin. Oncol., 31, 196, 2004. 

98. Sapra, P. and Allen, T. M., Ligand-targeted liposomal anticancer drugs, Prog. Lipid Res., 42, 439, 

2003. 

99. Carlsson, J. et al., Ligand liposomes and boron neutron capture therapy, J. Neurooncol., 62, 47, 2003. 

100. Hawthorne, M. F. and Shelly, K., Liposomes as drug delivery vehicles for boron agents, 

J. Neurooncol., 33, 53, 1997. 

101. Mishima, Y. et al., Treatment of malignant melanoma by single thermal neutron capture therapy with 

melanoma-seeking 10 B-compound, Lancet, 2, 388, 1989. 

102. Coderre, J. A. et al., Selective delivery of boron by the melanin precursor analogue p-boronophenylalanine 

to tumors other than melanoma, Cancer Res., 50, 138, 1990. 

103. Ono, K. et al., The combined effect of boronophenylalanine and borocaptate in boron neutron capture 

therapy for SCCVII tumors in mice, Int. J. Radiat. Oncol. Biol. Phys., 43, 431, 1999. 

104. Obayashi, S. et al., Delivery of (10)boron to oral squamous cell carcinoma using boronophenylalanine 

and borocaptate sodium for boron neutron capture therapy, Oral Oncol., 40, 474, 2004. 

105. Yoshino, K. et al., Improvement of solubility of p-boronophenylalanine by complex formation with 

monosaccharides, Strahlenther. Onkol., 165, 127, 1989. 

106. Ryynanen, P. M. et al., Models for estimation of the 10 B concentration after BPA-fructose complex 

infusion in patients during epithermal neutron irradiation in BNCT, Int. J. Radiat. Oncol. Biol. Phys., 

48, 1145, 2000. 

107. Pavanetto, F. et al., Boron-loaded liposomes in the treatment of hepatic metastases: preliminary 

investigation by autoradiography analysis, Drug Deliv., 7, 97, 2000. 

108. Martini, S. et al., Boronphenylalanine insertion in cationic liposomes for boron neutron capture 

therapy, Biophys. Chem., 111, 27, 2004. 

109. Smyth Templeton, N. et al., Cationic liposomes as in vivo delivery vehicles, Curr. Med. Chem., 10, 

1279, 2003. 

110. Mehta, S. C., Lai, J. C., and Lu, D. R., Liposomal formulations containing sodium mercaptoundecahydrododecaborate 

(BSH) for boron neutron capture therapy, J. Microencapsul., 13, 269, 1996. 

111. Ji, B. et al., Cell culture and animal studies for intracerebral delivery of borocaptate in liposomal 

formulation, Drug Deliv., 8, 13, 2001. 

112. Maruyama, K. et al., Intracellular targeting of sodium mercaptoundecahydrododecaborate (BSH) to 

solid tumors by transferrin-PEG liposomes, for boron neutron-capture therapy (BNCT), J. Control. 

Release, 98, 195, 2004. 

113. Valliant, J. F. et al., The medicinal chemistry of carboranes, Coord. Chem. Rev., 232, 173, 2002. 

114. Hawthorne, M. F., The role of chemistry in the development of cancer therapy by the boron-neutron 

capture reaction, Angew. Chem. Int. Ed. Engl., 105, 997, 1993. 

115. Moraes, A. M., Santanaand, M. H. A., and Carbonell, R. G., Preparation and characterization of 

liposomal systems entrapping the boronated compound o-carboranylpropylamine, 

J. Microencapsul., 16, 647, 1999. 



116. Moraes, A. M., Santana, M. H. A., and Carbonell, R. G., Characterization of liposomal systems 

entrapping boron-containing compounds in response to pH gradients, In Biofunctional Membranes, 

Proceedings of the International Conference on Biofunctional Membranes, Butterfield, A., Ed., 

Plenum Press, New York, p.259, 1996. 

117. Shelly, K. et al., Model studies directed toward the boron neutron-capture therapy of cancer: boron 

delivery to murine tumors with liposomes, Proc. Natl. Acad. Sci. USA, 89, 9039, 1992. 

118. Feakes, D. A. et al., [Na 3B 20H 17NH 3]: Synthesis and liposomal delivery to murine tumors, Proc. 

Natl. Acad. Sci. USA, 91, 3029, 1994. 

119. Hawthorne, M. F., Shelly, K., and Li, F., The versatile chemistry of the B 20H 18] 2K ions: novel 

reactions and structural motifs, Chem. Commun. (Camb.), 547, 2002. 

120. Feakes, D. A. et al., Synthesis and in vivo murine evaluation of [Na 41-(1 0 -B 10H 9)-6-SHB 10H 8]asa 

potential agent for boron neutron capture therapy, Proc. Natl Acad. Sci. USA, 96, 6406, 1999. 

121. Watson-Clark, R. A. et al., Model studies directed toward the application of boron neutron capture 

therapy to rheumatoid arthritis: boron delivery by liposomes in rat collagen-induced arthritis, Proc. 


122. Feakes, D. A., Spinler, J. K., and Harris, F. R., Synthesis of boron-containing cholesterol derivatives 

for incorporation into unilamellar liposomes and evaluation as potential agents for BNCT, 

Tetrahedron, 55, 11177, 1999. 

123. Thirumamagal, B. T. S., Zhao, X. B., Bandyopadhyaya, A. K., Narayanasamy, S., Johnsamuel, J., 

Tiwari, R., Golightly, D. W., Patel, V., Jehning, B. T., Backer, M. V., Barth, R. F., Lee, R. J., Backer, 

J. M., and Tjarks, W., Receptor-targeted liposomal delivery of boron-containing cholesterol mimics 

for boron neutron capture therapy (BNCT), Bioconjug. Chem., 2006, in press. 

124. Pan, G., Oie, S., and Lu, D. R., Uptake of the carborane derivative of cholesteryl ester by glioma 

cancer cells is mediated through LDL receptors, Pharm. Res., 21, 1257, 2004. 

125. Maletinska, L. et al., Human glioblastoma cell lines: levels of low-density lipoprotein receptor and 

low-density lipoprotein receptor-related protein, Cancer Res., 60, 2300, 2000. 

126. Nygren, C. et al., Increased levels of cholesterol esters in glioma tissue and surrounding areas of 

human brain, Br. J. Neurosurg., 11, 216, 1997. 

127. Leppala, J. et al., Accumulation of 99m Tc-low-density lipoprotein in human malignant glioma, Br. 

J. Cancer, 71, 383, 1995. 

128. Peacock, G. et al., In vitro uptake of a new cholesteryl carborane ester compound by human glioma 

cell lines, J. Pharm. Sci., 93, 13, 2004. 

129. Wei, Q., Kullberg, E. B., and Gedda, L., Trastuzumab-conjugated boron-containing liposomes for 

tumor-cell targeting; development and cellular studies, Int. J. Oncol., 23, 1159, 2003. 

130. Bohl Kullberg, E. et al., Development of EGF-conjugated liposomes for targeted delivery of boronated 

DNA-binding agents, Bioconjug. Chem., 13, 737, 2002. 

131. Stephenson, S. M. et al., Folate receptor-targeted liposomes as possible delivery vehicles for boron 

neutron capture therapy, Anticancer Res., 23, 3341, 2003. 

132. Yanagie, H. et al., Boron neutron capture therapy using 10 B entrapped anti-CEA immunoliposome, 

Hum. Cell, 2, 290, 1989. 

133. Yanagie, H. et al., Application of boronated anti-CEA immunoliposome to tumour cell growth 

inhibition in in vitro boron neutron capture therapy model, Br. J. Cancer, 63, 522, 1991. 

134. Yanagie, H. et al., Inhibition of growth of human breast cancer cells in culture by neutron capture 

using liposomes containing 10 B, Biomed. Pharmacother., 56, 93, 2002. 

135. Xu, L., Boron neutron capture therapy of human gastric cancer by boron-containing immunoliposomes 

under thermal neutron irradiation (in Chinese), Zhonghua Yi Xue Za Zhi., 71, 568, 1991. 

136. Xu, L., Zhang, X. Y., and Zhang, S. Y., In vitro and in vivo targeting therapy of immunoliposomes 

against human gastric cancer (in Chinese), Zhonghua Yi Xue Za Zhi., 74, 83, 1994. 

137. Pan, X. Q., Wang, H., and Lee, R. J., Boron delivery to a murine lung carcinoma using folate 

receptor-targeted liposomes, Anticancer Res., 22, 1629, 2002. 

138. Pan, X. Q. et al., Boron-containing folate receptor-targeted liposomes as potential delivery agents for 

neutron capture therapy, Bioconjug. Chem., 13, 435, 2002. 

139. Sudimack, J. J. et al., Folate receptor-mediated liposomal delivery of a lipophilic boron agent to 

tumor cells in vitro for neutron capture therapy, Pharm. Res., 19, 1502, 2002. 


102 


140. Bohl Kullberg, E. et al., Introductory experiments on ligand liposomes as delivery agents for boron 

neutron capture therapy, Int. J. Oncol., 23, 461, 2003. 

141. Kullberg, E. B., Nestor, M., and Gedda, L., Tumor-cell targeted epiderimal growth factor liposomes 

loaded with boronated acridine: uptake and processing, Pharm. Res., 20, 229, 2003. 

142. Mehvar, R., Dextrans for targeted and sustained delivery of therapeutic and imaging agents, 

J. Control. Release, 69, 1, 2000. 

143. Mehvar, R., Recent trends in the use of polysaccharides for improved delivery of therapeutic agents: 

pharmacokinetic and pharmacodynamic perspectives, Curr. Pharm. Biotechnol., 4, 283, 2003. 

144. Chau, Y., Tan, F. E., and Langer, R., Synthesis and characterization of dextran-peptide–methotrexate 

conjugates for tumor targeting via mediation by matrix metalloproteinase II and matrix metalloproteinase 

IX, Bioconjug. Chem., 15, 931, 2004. 

145. Zhang, X. and Mehvar, R., Dextran–methylprednisolone succinate as a prodrug of methylprednisolone: 

Plasma and tissue disposition, J. Pharm. Sci., 90, 2078, 2001. 

146. Larsson, B., Gabel, D., and Borner, H. G., Boron-loaded macromolecules in experimental physiology: 

tracing by neutron capture radiography, Phys. Med. Biol., 29, 361, 1984. 

147. Gabel, D. and Walczyna, R., B-Decachloro-o-carborane derivatives suitable for the preparation of 

boron-labeled biological macromolecules, Z. Naturforsch., C, 37, 1038, 1982. 

148. Pettersson, M. L. et al., In vitro immunological activity of a dextran-boronated monoclonal antibody, 

Strahlenther. Onkol., 165, 151, 1989. 

149. Ujeno, Y. et al., The enhancement of thermal-neutron induced cell death by 10-boron dextran, 

Strahlenther. Onkol., 165, 201, 1989. 

150. Pettersson, M. L. et al., Immunoreactivity of boronated antibodies, J. Immunol. Methods, 126, 95, 

1990. 

151. Holmberg, A. and Meurling, L., Preparation of sulfhydrylborane-dextran conjugates for boron 

neutron capture therapy, Bioconjug. Chem., 4, 570, 1993. 

152. Carlsson, J. et al., Strategy for boron neutron capture therapy against tumor cells with overexpression 

of the epidermal growth factor-receptor, Int. J. Radiat. Oncol. Biol. Phys., 30, 105, 1994. 

153. Gedda, L. et al., Development and in vitro studies of epidermal growth factor-dextran conjugates for 

boron neutron capture therapy, Bioconjug. Chem., 7, 584, 1996. 

154. Mehta, S. C. and Lu, D. R., Targeted drug delivery for boron neutron capture therapy, Pharm. Res., 

13, 344, 1996. 

155. Olsson, P. et al., Uptake of a boronated epidermal growth factor-dextran conjugate in CHO xenografts 

with and without human EGF-receptor expression, Anti-Cancer Drug Des., 13, 279, 1998. 

156. Novick, S. et al., Linkage of boronated polylysine to glycoside moieties of polyclonal antibody; 

boronated antibodies as potential delivery agents for neutron capture therapy, Nucl. Med. Biol., 29, 

159, 2002. 

157. Ferro, V. A., Morris, J. H., and Stimson, W. H., A novel method for boronating antibodies without 

loss of immunoreactivity, for use in neutron capture therapy, Drug Des. Discov., 13, 13, 1995. 

158. Sano, T., Boron-enriched streptavidin potentially useful as a component of boron carriers for neutron 

capture therapy of cancer, Bioconjug. Chem., 10, 905, 1999. 

159. Thomas, J. and Hawthorne, M. F., Dodeca(carboranyl)-substituted closomers: Toward unimolecular 

nanoparticles as delivery vehicles for BNCT, Chem. Commun. (Camb.), 1884, 2001. 

160. Pardridge, W. M., Vector-mediated drug delivery to the brain, Adv. Drug Deliv. Rev., 36, 299, 1999. 

161. Pardridge, W. M., Blood–brain barrier biology and methodology, J. Neurovirol., 5, 556, 1999. 

162. Hatakeyama, H. et al., Factors governing the in vivo tissue uptake of transferrin-coupled polyethylene 

glycol liposomes in vivo, Int. J. Pharm., 281, 25, 2004. 

163. Yanagie, H. et al., Accumulation of boron compounds to tumor with polyethylene–glycol binding 

liposome by using neutron capture autoradiography, Appl. Radiat. Isot., 61, 639, 2004. 

164. Maruyama, K. et al., Targetability of novel immunoliposomes modified with amphipathic poly 

(ethylene glycol)s conjugated at their distal terminals to monoclonal antibodies, Biochim. 

Biophys. Acta, 1234, 74, 1995. 

165. Yang, W. et al., Evaluation of systemically administered radiolabeled epidermal growth factor as a 

brain tumor targeting agent, J. Neurooncol., 55, 19, 2001. 

166. Bobo, R. H. et al., Convection-enhanced delivery of macromolecules in the brain, Proc. Natl Acad. 

Sci. USA, 91, 2076, 1994. 



167. Groothuis, D. R., The blood–brain and blood–tumor barriers: A review of strategies for increasing 

drug delivery, Neurooncology, 2, 45, 2000. 

168. Vogelbaum, M. A., Convection enhanced delivery for the treatment of malignant gliomas: symposium 

review, J. Neurooncol., 73, 57, 2005. 

169. Husain, S. R. and Puri, R. K., Interleukin-13 receptor-directed cytotoxin for malignant glioma 

therapy: from bench to bedside, J. Neurooncol., 65, 37, 2003. 

170. Kunwar, S., Convection enhanced delivery of IL13-PE38QQR for treatment of recurrent malignant 

glioma: presentation of interim findings from ongoing phase 1 studies, Acta Neurochir. Suppl., 88, 

105, 2003. 

171. Wikstrand, C. J. et al., Monoclonal antibodies against EGFRvIII are tumor specific and react with 

breast and lung carcinomas and malignant gliomas, Cancer Res., 55, 3140, 1995. 

172. Wikstrand, C. J. et al., Cell surface localization and density of the tumor-associated variant of the 

epidermal growth factor receptor, EGFRvIII, Cancer Res., 57, 4130, 1997. 

173. Barth, R. F., Wu, G., Kawabata, S., Sferra, T. J., Bandyopadhyaya, A. K., Tjarks, W., Ferketich, 

A. K., Binns, P. J., Riley, K. J., Coderre, J. A., Ciesielski, M. J., Fenstermaker, R. A., Wikstrand, C. J. 

et al., Molecular targeting and treatment of EGFRvIII positive gliomas using boronated monoclonal 

antibody L8A4, Clin. Cancer Res., 12, 3792, 2006. 

174. Cokgor, I. et al., Phase I trial results of iodine-131-labeled antitenascin monoclonal antibody 81C6 

treatment of patients with newly diagnosed malignant gliomas, J. Clin. Oncol., 18, 3862, 2000. 

175. Akabani, G. et al., Dosimetry and radiographic analysis of 131I-labeled anti-tenascin 81C6 murine 

monoclonal antibody in newly diagnosed patients with malignant gliomas: a phase II study, J. Nucl. 

Med., 46, 1042, 2005. 

176. Laske, D. W., Youle, R. J., and Oldfield, E. H., Tumor regression with regional distribution of the 

targeted toxin TF-CRM107 in patients with malignant brain tumors, Nat. Med., 3, 1362, 1997. 

177. Sampson, J. H. et al., Progress report of a Phase I study of the intracerebral microinfusion of a 

recombinant chimeric protein composed of transforming growth factor (TGF)-alpha and a mutated 

form of the Pseudomonas exotoxin termed PE-38 (TP-38) for the treatment of malignant brain 

tumors, J. Neurooncol., 65, 27, 2003. 

178. Weber, F. et al., Safety, tolerability, and tumor response of IL4-Pseudomonas exotoxin (NBI-3001) 

in patients with recurrent malignant glioma, J. Neurooncol., 64, 125, 2003. 

179. Weber, F. W. et al., Local convection enhanced delivery of IL4-Pseudomonas exotoxin (NBI-3001) 

for treatment of patients with recurrent malignant glioma, Acta Neurochir. Suppl., 88, 93, 2003. 

180. Ferrari, M., Nanovector therapeutics, Curr. Opin. Chem. Biol., 9, 343, 2005. 

181. Ferrari, M., Cancer nanotechnology: opportunities and challenges, Nat. Rev. Cancer, 5, 161, 2005. 

182. Blue, T. E. and Yanch, J. C., Accelerator-based epithermal neutron sources for boron neutrom 

capture therapy of brain tumors, J. Neurooncol., 62, 19, 2003. 


7 

CONTENTS 

Preclinical Characterization of 

Engineered Nanoparticles 

Intended for Cancer 

Therapeutics 

Anil K. Patri, Marina A. Dobrovolskaia, Stephan T. Stern, and 

Scott E. McNeil 

7.1 Introduction ......................................................................................................................... 106 

7.1.1 Physicochemical Characterization .......................................................................... 106 

7.1.2 In Vitro Characterization ........................................................................................ 107 

7.1.3 In Vivo Characterization ......................................................................................... 107 

7.2 Physicochemical Characterization ...................................................................................... 107 

7.2.1 Physicochemical Characterization Strategies ......................................................... 108 

7.2.2 Instrumentation........................................................................................................ 108 

7.2.2.1 Spectroscopy ............................................................................................ 108 

7.2.2.2 Chromatography ....................................................................................... 109 

7.2.2.3 Microscopy ............................................................................................... 109 

7.2.2.4 Size and Size Distribution ....................................................................... 110 

7.2.2.5 Surface Characteristics............................................................................. 112 

7.2.2.6 Functionality............................................................................................. 112 

7.2.2.7 Composition and Purity ........................................................................... 113 

7.2.2.8 Stability .................................................................................................... 114 

7.3 In Vitro Pharmacological and Toxicological Assessments................................................ 115 

7.3.1 Special Considerations ............................................................................................ 115 

7.3.2 In Vitro Target-Organ Toxicity .............................................................................. 116 

7.3.3 Cytotoxicity ............................................................................................................. 117 

7.3.4 Oxidative Stress....................................................................................................... 118 

7.3.5 Apoptosis and Mitochondrial Dysfunction ............................................................. 119 

7.3.6 Proteomics and Toxicogenomics ............................................................................ 121 

7.4 In Vivo Pharmacokinetic and Toxicological Assessments ................................................ 121 

7.5 Immunotoxicity ................................................................................................................... 124 

This project has been funded in whole or in part with federal funds from the National Cancer Institute, National Institutes of 

Health, under contract N01-CO-12400. The content of this publication does not necessarily reflect the views or policies of the 

Department of Health and Human Services, nor does mention of trade names, commercial products, or organizations imply 

endorsement by the U.S. Government. 


105

106 

7.5.1 Applicability of Standard Immunological Methods for Nanoparticle Evaluation 

and Challenges Specific to Nanoparticle Characterization .................................... 125 

7.5.1.1 Blood Contact Properties ......................................................................... 125 

7.5.2 Immunogenicity....................................................................................................... 128 

7.6 Conclusions ......................................................................................................................... 128 

References .................................................................................................................................... 129 


As discussed in other chapters of this book, nanotechnology offers the research community the 

potential to significantly transform cancer diagnostics and therapeutics. Our ability to manipulate 

the biological and physicochemical properties of nanomaterial allows for more efficient drug 

targeting and delivery, resulting in greater potency and specificity, and decreased adverse side 

effects. The combinatorial possibilities of these multifunctional platforms have been the focus of 

considerable research and funding, but realization of their use in clinical trials is highly dependent 

on rigorous preclinical characterization to meet regulatory provisions and elucidated structure– 

activity relationships (SARs). A rational characterization strategy for biomedical nanoparticles 

contains three essential components (see Figure 7.1): physicochemical characterization, in vitro 

assays, and in vivo studies. 

7.1.1 PHYSICOCHEMICAL CHARACTERIZATION 

One of the major criticisms of early biomedical nanotechnology research was the general lack of 

physicochemical characterization that did not allow for the meaningful interpretation of resulting 

data or inter-laboratory comparisons. Traditional small molecule drugs are characterized by data that 

contribute to the chemistry, manufacturing and controls (CMC) section of the investigational new 

drug (IND) application with Food and Drug Administration (FDA), which include their molecular 

weight, chemical composition, identity, purity, solubility, and stability. The instrumentation to 

ascertain these properties has been well established, and the techniques are standardized. Engineered 

nanomaterials have dimensions between small molecules and bulk materials and often exhibit 

different physical and chemical properties than their counterparts. These physical and chemical 

Physical 

Characterization: 

Size 

Shape 

Composition 

Molecular weight 

Surface chemistry 

Identity 

Purity 

Stability 

Solubility 

In Vitro: 

Pharmacology 

Blood contact properties 

Effects on immune cell 

function 

Cytotoxicity 

Mechanistic toxicology 

Sterility 

In Vivo: 

ADME 

Safety 

Efficacy 

FIGURE 7.1 An assay cascade for preclinical characterization of nanomaterials. 



Preclinical Characterization of Engineered Nanoparticles Intended for Cancer Therapeutics 107 

properties influence the biological activity of nanoparticles and may depend on parameters such as 

particle size, size distribution, surface area, surface charge, surface functionality, shape, and aggregation 

state. Additionally, because most nanoparticle concepts are multifunctional, the 

distribution of targeting, imaging, and therapeutic components can also have dramatic effects on 

nanoparticle biological activity. There is a need to establish and standardize techniques to define 

these nanoparticle attributes. 

There is now ample evidence that size and surface characteristics can dramatically affect 

nanoparticle behavior in biological systems. 1–4 For instance, a decrease in particle size leads to 

an exponential increase in surface area per unit mass, and an attendant increase in the availability of 

reactive groups on the surface. Nanoparticles with cationic surface character have a notably 

increased ability to cross the blood–brain barrier compared to nanoparticles with anionic surfaces. 5 

In general, surface area, rather than mass, provides a better fit of dose–response relationships in 

toxicity studies for particles of various sizes. 6,7 Physicochemical characterization of properties, 

such as size, surface area, surface chemistry, and aggregation/agglomeration state, can provide the 

basis for better understanding of SARs. 

7.1.2 IN VITRO CHARACTERIZATION 

In vitro assays enable the isolation and analysis of specific biological and mechanistic pathways 

under controlled conditions. While many in vitro assays for nanomaterials will be similar to those 

used for traditional drugs, others will address mechanisms more specific to nanoparticles, such as 

oxidative stress. Noncellular assays measuring processes, such as protein adsorption, will also be an 

important accompaniment to cell-based assays. For example, monitoring the profile of serum 

proteins that absorb to nanoparticles in an in vitro environment may further our understanding of 

how nanoparticles interact with components of the reticuloendothelial system (RES) in vivo. 8,9 

Additionally, proteomics and toxicogenomics can be employed to potentially identify biomarkers 

of toxicity related to nanomaterial exposure. 10 

7.1.3 IN VIVO CHARACTERIZATION 

In vivo studies must be conducted to better understand the safety and behavior of nanoparticles in a 

living organism. As with any new chemical entity (NCE), the nanoparticle formulations’ pharmacological 

and toxicological properties (i.e., ADME/Tox) need to be thoroughly characterized. 

In vivo studies should include examination of nanoparticles’ effects on various organs and 

systems, such as the liver, heart, kidney, and immune system. 

In this chapter, we outline the scientific rationale underlying the development of an assay 

cascade, with special attention paid to the selection and adaptations of assays and analytical protocols 

needed to extract meaningful efficacy and safety data from nanomaterials. These are presented in the 

following four sections: (1) physicochemical characterization, (2) in vitro pharmacology and toxicology 

assessment, (3) in vivo pharmacology and toxicology assessment, and (4) 

immunotoxicity. Standardized characterization of nanomaterials will facilitate better inter-laboratory 

comparison of results and will enhance the quality of scientific data submitted by investigators in 

support of their IND applications. 

7.2 PHYSICOCHEMICAL CHARACTERIZATION 

The physicochemical characteristics of nanomaterials can affect their cellular uptake, binding to 

blood proteins, access to target sites, and ability to cause damage to cells and tissue. 11 Standard 

methods for physicochemical characterization of nanomaterials will provide the basis for rational 

product development as well as consistent and interpretable results for tests of efficacy and safety. 

Few examples of standard characterization criteria exist in the literature, and there is as yet no 


108 

consensus as to what measurement criteria are appropriate for any given nanomaterial product. 

However, it is clear that the diversity and complexity of nanomaterials used in biomedical applications 

dictates a more comprehensive and strategic approach to characterization than has been 

applied to date. 

There are many varieties of nanomaterials currently being investigated for biomedical applications, 

especially for cancer diagnosis, imaging, and targeted drug delivery. These nanomaterials 

may be classified under several broad categories: 

† Organic nanoparticles (e.g., dendrimers, polymers, functionalized fullerenes) 

† Inorganic nanoparticles and organic–inorganic hybrids (e.g., iron oxide core particles, 

quantum dots) 

† Liposomes and other biological nanomaterials 

Each category of nanomaterial has a distinctly different composition that gives rise to different 

physical properties, such as solubility, stability, surface characteristics, and functional capabilities. 

Even within a single category, there can be a tremendous variety of product compositions, each 

with unique physical and chemical characteristics, and each requiring a different strategy for 

measuring those characteristics. This section examines the various categories of nanoparticles, 

and the tools and instrumentation available to address physicochemical characterization. 

7.2.1 PHYSICOCHEMICAL CHARACTERIZATION STRATEGIES 

Successful characterization strategies will enable one to begin associating the physicochemical 

properties of a nanomaterial with its in vivo behavior (i.e., SARs). This is an important step in the 

development of any material used for medical applications. For small molecules, the basis of most 

traditional drugs, the characterization techniques have been well established and standardized to 

determine their attributes, such as melting point, boiling point, molecular weight and structure, 

identity, composition, solubility, purity, and stability. These characteristics are measured and 

adequately defined using elemental analysis, mass spectrometry (MS), nuclear magnetic resonance 

(NMR), ultraviolet-visible (UV–vis) spectrophotometry, infrared (IR) spectroscopy, high-performance 

liquid chromatography (HPLC), gas chromatography (GC), capillary electrophoresis 

(CE), polarimetry, and other common analytical methods. Each of these individual techniques 

provides unique information about the sample, while together they provide the foundation for 

product quality control, manufacturing, and regulatory approval. 

Many of the techniques used to characterize small molecules apply to nanomaterials. However, 

due to the composite nature of nanomaterials, the definition and measurement of these attributes can 

be quite different. To fully understand the attributes of a nanomaterial, additional characterizations 

are needed, such as size, surface chemistry, surface area, polydispersity, and zeta potential (see 

Figure 7.2). A comprehensive analysis of these properties is necessary to better understand in vivo 

effects and to allow for greater consistency and reproducibility in their preparation. The requirements 

set by regulatory bodies for quality control and consistency of biomedical nanomaterials are 

likely to be as stringent as those for small molecule preparations, but the path to verifying quality 

will require a more sophisticated approach. At the core of this analysis is an array of tools and 

instrumentation that are particularly well suited to measuring the properties of nanomaterials. 

7.2.2 INSTRUMENTATION 

7.2.2.1 Spectroscopy 


Many traditional analytical methods can be applied to the characterization of nanomaterials. For 

example, NMR is extensively used to characterize dendrimers, polymers, and fullerenes 

derivatives, and provides unique information on the structure, purity, and functionality. 12–14 



O O 

N 

H OH 

HO 

O 

O 

In addition, the average number of terminal capping groups, number of small molecule ligands, and 

drugs in a multifunctional nanomaterial can be ascertained by comparing the integration values 

with chemical shifts unique to the ligands. UV–vis absorption spectrophotometry is also extensively 

used to identify and quantify the chromophore present in the preparation by using its 

extinction coefficient. Spectrofluorimetry is used in cases where the material has inherent fluorescence 

(such as quantum dots) or labeled with a fluorescence probe. Matrix-assisted laser 

desorption ionization time-of-flight (MALDI-TOF) MS is used extensively for macromolecules, 

dendrimers, and polymers to determine the molecular weight and utilize novel matrices to minimize 

the fragmentation of the macromolecule before reaching the detector. In the case of lower generation 

dendrimers, the presence of impurities, incomplete reaction, and reaction byproducts can be 

easily determined using MS. 

7.2.2.2 Chromatography 

O 

Small molecules 

• Elemental analysis 

• Mass spectrometry 

• NMR spectroscopy 

• UV−vis spectroscopy 

• IR spectroscopy 

• HPLC 

• GC 

• Polarimetry 

O 

Liquid chromatography methods such as analytical HPLC and size-exclusion chromatography 

(SEC; also called gel-permeation chromatography or GPC) utilize a column to separate components 

of a mixture in a liquid mobile phase based on their interaction with a solid stationary phase. The 

eluents are passed through UV–vis and fluorescence detectors with a flow cell where the absorbance 

and fluorescence is recorded to determine the purity of the sample. Although these techniques are 

suitable for stable polymers, dendrimers, 15 functionalized fullerenes, and protein- and peptide-based 

nanomaterial, they are not suitable for particles that degrade under experimental conditions or have 

excessive nonspecific binding to the solid matrix. 

7.2.2.3 Microscopy 

O OH 

H 

O 

O 

–H 

O 

Physicochemical Parameters 

• Composition 

• Physical properties 

• Chemical properties 

• Identification 

• Quality 

• Purity 

• Stability 

Liposome 

Dendrimer Gold nanoshell 

Quantum dot Fullerene 

Nanomaterial 

• Microscopy (AFM, TEM, SEM) 

• Scattering (DLS, MALLS, SAXS, 

SANS), FCS, AU 

• SEC, FFF 

• Electrophoresis (CE, PAGE) 

• Zeta potential 

• Fluorimetry 

FIGURE 7.2 Physicochemical characterization methods and instrumentation for small molecules and 

nanotechnology. 

Scanning probe microscopy (SPM) techniques can be employed to measure the size, topography, 

composition, and structural properties of nanoparticles. Related techniques such as scanning 

tunneling microscopy (STM), electric field gradient microscopy (EFM), scanning thermal 

microscopy, and magnetic field microscopy (MFM) combined with atomic force microscopy 

(AFM), can be used to investigate the structural, electronic, thermal, and magnetic properties of 


110 

a nanomaterial. AFM uses a nanoscale probe to detect the inter-atomic forces and interactions 

between the probe and the material being analyzed and is capable of determining size and shape 

within a spatial resolution of a few angstroms. 16 Apart from the ability to measure the particle size 

in a dry state as well as in aqueous and physiological conditions, AFM is a useful tool to probe the 

interaction of nanoparticles with supported lipid bilayers. This technique has been successfully 

used to compare nanoparticle interactions in in vitro cell assays. 17,18 The ability to image under 

physiological conditions makes AFM a powerful tool for the characterization of nanoparticles in a 

dynamic, biological context. A variant of this method, molecular recognition force microscopy 

(MRFM), can be employed to study the specific ligand–receptor interactions between nanoparticles 

and their biological targets. 

Optical microscopy techniques are useful at the micron scale and are extensively used for 

imaging structural features. Fluorescence and confocal microscopy may be used to determine 

cellular binding and internalization of fluorescent-labeled nanoparticles 19 or those that are inherently 

fluorescent, such as quantum dots. But a more precise analysis of nanomaterial size and other 

direct measurements of physical properties will require a more sophisticated and specialized set of 

microscopic and spectroscopic techniques. 

Scanning electron microscopy (SEM) provides information on the size, size distribution, shape, 

and density of nanomaterials. Transmission electron microscopy (TEM) and high-resolution TEM 

are more powerful than SEM in providing details at the atomic scale and can yield information 

regarding the crystal structure, quality, and grain size. TEM can be coupled with other characterization 

tools, such as electron energy loss spectrometry (EELS) or energy dispersive x-ray 

spectrometry (EDS), to provide additional information on the electronic structure and elemental 

composition of nanomaterials. Samples for TEM are evaluated dry or in a frozen state, under highvacuum 

conditions. Nanoparticles analyzed by this instrument must therefore be stable under these 

extreme conditions. Additionally, while considered a gold standard of microscopic characterization 

methods, TEM requires a great deal of skill and time to obtain good data. In principle, when 

establishing characterization protocols, TEM can be used to validate characterization methods 

that are easier to use on a routine basis. Further description of analytical technologies as they 

apply to the measurement of specific nanomaterial properties is provided in the following sections. 

7.2.2.4 Size and Size Distribution 


Size is one of the critical parameters that dictate the absorption, biodistribution, and route of 

elimination for biomedical nanomaterials. 20 Generally, nanoparticles with dimensions of less 

than 5–10 nm are rapidly cleared after systemic administration, while particles from 10 to 70 nm 

in diameter may penetrate capillary walls throughout the body. 21,22 Larger particles 70–200 nm 

often remain in circulation for extended times. 22,23 This general correlation of biodistribution and 

elimination with respect to size may vary greatly depending on nanoparticle surface characteristics. 

Specifically in cancer applications, size is an important factor in the accumulation of therapeutic 

nanomaterials in tumors, usually as a result of enhanced permeation and retention (EPR), 

caused by local defects in the vasculature and poor lymphatic drainage. 24 Particle size can be 

precisely tuned to take advantage of this phenomenon and passively target and deliver a therapeutic 

payload to tumors. 25–27 

Depending on the category of the nanomaterial, synthesis and scale-up can be problematic. 

Most biomedical nanomaterials for therapeutic and diagnostic applications are complex and 

involve some combination of molecular self-assembly, encapsulation, and/or the use of nanosized 

metal or polymer cores, surfactants and/or proteins to impart solubility and functionality. 

Due to inherent variability in the manufacturing process, one rarely achieves a monodisperse, 

homogeneous product. It is therefore important to ascertain the precise size, size distribution, 

and polydispersity index (PDI) of the material. There are several techniques available to assess 

these parameters, including electron microscopy, AFM, and light scattering. Light scattering 



techniques can measure overall size and polydispersity of the particles. TEM is powerful in 

ascertaining the homogeneity of nanoparticles with encapsulated metals and in determining core 

size. With knowledge of nanoparticle geometry and size, surface area can also be estimated. 

For biological applications, it is important to measure the physical characteristics of the nanomaterial 

in isotonic solution at physiological pH and temperature. The hydrodynamic size can be 

measured under these conditions using dynamic light scattering (DLS) (also known as photon 

correlation spectroscopy [PCS] and quasi elastic light scattering [QELS]) and analytical ultracentrifugation 

(AU). In a DLS experiment, the effects of Brownian motion (particle movement caused 

by random collisions in solution) provide information on particle size and size distribution. The 

sample is illuminated with a laser, and the intensity fluctuations in the scattered light are analyzed 

and related to the size of the suspended particles. This technique is useful in determining whether 

the nanomaterial is monodisperse in size distribution. These data are influenced by the viscosity and 

the temperature of the medium, since Brownian motion depends on these factors. The pH of the 

medium and salt concentration may also affect the degree of agglomeration in some samples. With 

DLS, sample preparation is easy, the measurement is quick, and data are reproducible on larger 

sample volumes compared to microscopy techniques; however, better standardization of 

procedures, conditions, and data analysis tools will be required. Static light scattering provides 

information on molar mass and root-mean-squared (rms) radius for fractionated or monodisperse 

samples. One limitation of light scattering instruments is the inability to measure the size when the 

nanoparticles absorbs in the wavelength of the laser being used. Small-angle x-ray scattering 

(SAXS) 28 and small-angle neutron scattering (SANS) 29 can be used to measure the size, shape 

and orientation of components. Due to their cost and infrastructure requirements, there is limited 

availability of these instruments. For fluorescent nanomaterials such as quantum dots, size can be 

measured using fluorescence correlation spectroscopy (FCS). 30 

The hydrodynamic size of nanoparticles can also be measured with AU, which is traditionally 

used to measure the size of proteins. 31 The instrument spins the protein sample solution under high 

vacuum at a controlled speed and temperature while recording concentration distribution at set 

times. Even though this technique is designed to measure the size of proteins in solution, it has 

potential applications in the measurement of the hydrodynamic size of nanoparticles samples that 

are stable under the experimental conditions. Fractionation using SEC separates stable polymers 

into individual components and helps in the determination of the PDI. In the case of unfractionated 

samples, batch mode measurement provides averaged quantities such as weight-averaged molar 

mass and z-average rms radius. This technique is especially useful when combined with a refractive 

index detector to obtain absolute molecular weight for very high molecular weight polymers where 

traditional MS methods fail. 

In cases where the separation and fractionation of nanomaterial is not possible using a 

column with a stationary phase, such as when the nanomaterial may interact with the column 

packing material and render it unstable, asymmetric-flow field flow fractionation (AFFF) is 

useful. 32 In AFFF, separation occurs when the sample passes through a narrow channel with 

a cross-flow through a porous semi-permeable membrane. The faster moving smaller particles 

rise to the top of the flow and come out first followed by larger particles that stay closer to the 

membrane and migrate more slowly. One advantage in this method is that there is no stationary 

phase in the separation: the sample injected comes out intact with little loss of material due to 

nonspecific binding. This feature is particularly useful for less stable nanoparticles such as 

liposomes, or for polymer- or protein-coated metal nanoparticles that would otherwise interfere 

with the performance of a traditional GPC column. The efficiency of separation for AFFF is not 

as good as with GPC, but there have been recent improvements in instrumentation that are 

closing the gap in performance. For both GPC and AFFF, the quantity and hydrodynamic size of 

the nanoparticles are detected in eluted peaks by measuring absorbance, refractive index, and 

light scattering. 


112 

In addition to size, the shape of a nanoparticle may affect its distribution and absorption in the 

body. Spherical, tubular, plate-like, or nano-porous materials of the same composition can vary 

significantly in their surface energy, biological activity, and access to different physiological 

structures, such as cell walls, capillary vessels, etc. Methods such as AFM, SEM, TEM, and 

STM can be used to determine the distribution of shape in a nanoparticle preparation. 

7.2.2.5 Surface Characteristics 

Surface characteristics contribute to the nanoparticle’s solubility, aggregation tendency, ability to 

traverse biological barriers (such as a cell wall), biocompatibility, and targeting ability. The 

nanoparticle surface is also responsible for interaction and binding with plasma proteins in vivo, 

which in turn may alter the nanoparticle’s distribution and pharmacokinetics. For multifunctional 

nanoparticles, modifying agents are often attached to the surface to bind to receptors in target 

tissues and organs. The presence of charged functionalities on the nanoparticle surface may 

increase nonspecific uptake, making the preparation less effective in targeting. It has been shown 

that dendrimer nanoparticles displaying positively charged amine groups on their surface can be 

significantly more hemolytic and cytotoxic than nanoparticles displaying negatively charged 

carboxylates. 20 The negatively charged nanoparticles were also cleared more slowly from the 

blood compared to positively charged species, following intravenous administration to rats. 20 

Another potential effect of surface charge is to alter a nanoparticle’s ability to penetrate the 

blood–brain barrier. Studies have shown that for emulsifying wax nanoparticles, anionic surfaces 

were superior to neutral or cationic surfaces for penetration of the blood–brain barrier. 33 

Surface characteristics can be tuned to improve receptor binding, reduce toxicity, or alter 

biodistribution. For example, when the above-mentioned dendrimers were acetylated to neutralize 

exposed surface charges, the toxic effects of the nanoparticles were also neutralized. 20,34 Surface 

properties can also lead to toxicity through interaction with molecular oxygen, leading to oxidative 

stress and inflammation. Electron capture at the surface of the nanoparticle results in the formation 

of the superoxide radical, which can set off a cascade of reactions (e.g., through Fenton reaction or 

disumation) to generate reactive oxygen species (ROS). ROS generation has been studied extensively 

for inhaled nanoparticles, 11,35 and has been observed in engineered nanoparticles such as 

fullerenes, single walled nanotubes (SWNTs), and quantum dots. 6,36–44 Studies have shown a direct 

correlation between nanoparticle surface area and ROS-generating capacity and inflammatory 

effects. 11 

The nature and integrity of nanomaterial surfaces must be established through analytical 

measurements to ensure product quality and account for surface-dependent effects on biodistribution 

and toxicity. Potentiometric titrations provide crucial information on the net charge of a 

nanoparticle, and include zeta potential analysis, which provides information on the net charge 

and distribution under physiological conditions. Polyacrylamide gel electrophoresis (PAGE) 

analysis of dendrimers and other nanopolymers yields information on the molecular weight and 

the polydispersity of nanoparticles (such as trailing generations in dendrimer populations) based on 

their migration through the gel under an electric field. PAGE is also a powerful tool in the 

qualitative analysis of bioconjugates of nanomaterials with DNA, oligonucleotides, antibodies, 

and other ligands. Further analysis of the surface charge distribution and polydispersity of nanomaterials 

can be conducted using CE. MS is also effective in ascertaining the number and 

distribution of charges, especially for smaller and purer nanoparticles with known 

molecular weight. 

7.2.2.6 Functionality 


Analysis of the functional components of nanomaterials, such as targeting, imaging, and 

therapeutic agents, is critical to understand the in vivo efficacy of the preparation. Characteristic 



features of functional components include their quantity, distribution, orientation, and activity. For 

targeting agents, a key advantage of their use in nanoparticles is their ability to provide increased 

avidity to the target due to polyvalency. The level of polyvalency and activity of targeting agents 

can be monitored using surface plasmon resonance (SPR) to measure the rate constants for nanoparticle 

association and dissociation. During preclinical development, the affinity of nanomaterial 

preparations for their target molecule/receptor can be analyzed using SPR and compared to data 

obtained for binding to cellular receptors in culture. 45 

The average number of targeting agents per nanoparticle has to be optimized for both solubility 

and binding affinity. Affinity chromatography or SEC can be employed with some nanoparticles to 

separate nanoparticles with targeting agents from those without targeting agents. In nanoparticles 

containing antibodies 19,46 or proteins, quantification can be achieved using an enzyme-linked 

immunosorbent assay (ELISA) or bicinchoninic acid assay (BCA) if the inherent property of the 

nanoparticle itself does not interfere with the assay. In the case of dendrimers, NMR has been 

successfully applied to analyze the average number of targeting agents by comparing the integration 

values of the signals associated with the targeting agents to those belonging to the 

dendrimer. This is still an averaged technique that cannot distinguish the distribution of targeting 

agent density on a population of nanoparticles. 

For targeted drug delivery applications, it is obviously important for both the targeting and 

therapeutic agents to be on the same particle. If the therapeutic has UV–vis absorption, it can be 

quantified using UV–vis spectroscopy with the extinction coefficient of the drug. HPLC analysis is 

possible in some cases to evaluate the amount of the drug present in a known amount of material, 

after isolating the drug from the sample. 

7.2.2.7 Composition and Purity 

Biomedical nanomaterials can be comprised of a wide variety of substances, including polymers, 

metals and metal oxides, lipids and other organic compounds, and large biomolecules such as 

protein or DNA. In most cases, the nanomaterials combine two or more of these substances, 

such as in a core or shell of a particle, and in encapsulated or conjugated material. Analysis of 

chemical composition will be critical for confirming the purity and homogeneity of nanomaterial 

product preparations. 

Elemental analysis, such as CHN analysis, is most often used to ascertain the purity of small 

molecules. For nanomaterials, elemental analysis can be used to determine the composition and 

ratios of different elements present in the sample. For example, this technique can be used to 

determine the amount of linker present, if a unique element (such as sulfur) has been employed 

in the synthesis. In the case of core–shell metal nanoparticles, the ratio of core to shell material 

ratios can be determined. 

Atomic absorption (AA) and atomic emission (AE) spectroscopies can also be utilized to 

determine the composition of nanomaterials. For imaging applications using iron oxide nanoparticles 

or gadolinium (Gd)-based chelates, composition analysis is very important to quantify metals 

present in the preparation which influence imaging efficacy. Inductively coupled plasmon optical 

emission spectroscopy (ICP-OES) is very sensitive to determine the amount of Gd in such contrast 

agent conjugates. 47 Specific T1/T2 relaxivities of magnetic resonance contrast agents can, of 

course, be assessed under in vitro conditions in the actual MRI instrument. 

The purity of synthetic small molecules can be determined with a high degree of certainty since 

the analyte usually consists of a single component. With nanomaterials, purity must be determined 

in the context of multiple layered, conjugated, and encapsulated components. Purity analysis must 

account for the presence of solvents, free metals and chelates, unconjugated therapeutic or other 

agents, precursors, dimers, etc., that result in artifacts and side products of the preparation. 48 

Characterization of the inhomogeneity in ligand distribution is very important for efficacy as 

well as testing batch-to-batch reproducibility. 49 Proper methods and techniques to detect the 


114 

presence of all these entities are required to ensure the purity and quality of nanomaterial preparations 

and to further expand our understanding of SARs. 

7.2.2.8 Stability 


The ability of multifunctional nanoparticles to combine targeting, therapeutic, and imaging modalities 

is a key aspect of their versatility and anticipated clinical impact. 50–52 With such complex 

compositions, the stability of all the components in nanoparticles is essential to their biological 

function. Premature release of any of the components from the composite preparation may render it 

ineffective. For example, in a nanodelivery system containing a targeting agent and a drug, the 

nanoparticles with the drug cannot bind to the desired targeting site if the targeting agent is 

prematurely cleaved or released. If the drug is prematurely released, even if the nanoparticle 

reaches its target, there will no longer be a therapeutic benefit. 53 For this reason, it is important 

to determine the in vitro functional component stability under physiological conditions. 

For a nanomaterial providing targeted or timed-release drug delivery with an encapsulated 

drug, the release profile should be determined at different ionic strength, pH, and temperature 

conditions. Examples of such conditions include the stability at pH 7.4, in buffers such as phosphate 

buffered saline (PBS), and serum at 378C. There are many nanoparticle designs being pursued 

which incorporate the selective release of components triggered by an external stimulus after 

targeted delivery. If a therapeutic attached to a nanoparticle uses a cleavable linkage, the efficiency 

of release should be determined under the expected cleavage conditions. 54 

In cases where a metal complex is used (for example, a Gd chelate for enhanced MRI contrast), 

the stability constants for the encapsulation or complexation should be determined, since any 

release of free heavy metal will increase the in vivo toxicity of the preparation. 55 The potential 

in vivo application of quantum dots has raised some concerns that the CdSe core might be exposed 

by the breakdown of its protective polymer or inorganic shell, releasing the highly toxic heavy 

metal Cd 2C ions into the bloodstream. 56 The quantum dot shells have been designed to be protective, 

but their long-term stability (e.g., susceptibility to Cd leaching) has not been established. 

Studies conducted on primary hepatocytes in vitro suggest that CdSe core quantum dots may be 

acutely toxic under certain conditions. 57 Other studies suggest that under physiological conditions, 

appropriately coated quantum dots do not expose the host organism to toxic levels of the core 

material. 58–60 Apparently conflicting evidence as to the safety of quantum dots highlights the 

necessity of clearly and objectively establishing the stability of these nanoparticles under physiological 

conditions using standardized methodologies. 

It is also important to determine the stability of the nanoparticle under nonphysiological 

conditions to account for the effects of short-term and long-term storage, lyophilization, ultrafiltration, 

thermal exposure, pH variation, freeze–thawing, and exposure to light. 

In summary, adequate physicochemical characterization of nanomaterials should be included 

as an essential requirement for preclinical characterization. Just as molecular characterization 

forms the basis of dosing and toxicity studies for small molecule therapeutics and diagnostic 

compounds, physicochemical characterization provides the foundation for dosing and toxicity 

studies for nanomaterials intended for clinical applications. Standardized protocols are being established 

by Standards Developing Organizations, such as the International Standards Organization 

(ISO) and American Society for Testing and Materials (ASTM), for characterizing the many types 

of biomedical nanomaterials being developed today for human use. Additionally, standardized 

reference material (SRM) will enable analytical technologies to be calibrated and protocols to be 

tested for consistency and to facilitate inter-laboratory comparisons. 

To better control for the results of in vivo studies of nanomaterial absorption, distribution, 

metabolism, elimination, and toxicity, it will be necessary to examine the material in the same 

physicochemical state as would be found under physiological conditions. Particle-specific attributes 

that should be evaluated include surface characteristics, chemical composition, shape, size, 



and ligand dispersity. Additional properties that are influenced by experimental conditions include 

solubility, stability, protein binding, and aggregation state. Knowing the exact physiological conditions 

in different tissues and organs and developing a means to either replicate those conditions or 

measure physicochemical properties in situ is a significant challenge. But continued studies in this 

area will provide further data to elucidate the linkages between physicochemical characteristics of 

nanomaterials and their biological effects (i.e., SARs). 

7.3 IN VITRO PHARMACOLOGICAL AND TOXICOLOGICAL ASSESSMENTS 

Prior to filing an IND or investigational device exemption (IDE) application with the FDA and 

subsequent clinical testing in humans, a new product must be adequately studied for efficacy and 

safety using animal models. The cost- and labor-intensiveness of these in vivo studies impel drug 

and device researchers to make use of predictive in vitro methodologies wherever technology 

permits. In vitro models can serve as an initial assessment of a nanomaterial’s efficacy and absorption, 

distribution, metabolism, elimination, and toxicity (ADME/Tox), allowing a more strategic 

approach to animal studies. Used iteratively with in vivo studies, the two approaches can inform 

each other and help narrow investigations of the physiological and biochemical pathways that 

contribute to ADME/Tox behavior. 

A variety of cell-based in vitro systems are available, including perfused organs, tissue slices, 

cell cultures based on a single cell line or combination of cell lines, and primary cell preparations 

freshly derived from organ and tissue sources. In vitro models allow examination of biochemical 

mechanisms under controlled conditions, including specific toxicological pathways that may occur 

in target organs and tissues. Examples of mechanistic toxicological endpoints assessed in vitro 

include inhibition of protein synthesis and microtubule injury. These mechanistic endpoints can 

provide information not only as to the potential mechanisms of cell death, but also can identify 

compounds that may cause chronic toxicities that often results from sublethal mechanisms that may 

not cause overt toxicity in cytotoxicity assays. Common mechanistic paradigms associated with 

nanoparticle toxicity include oxidative stress, apoptosis, and mitochondrial dysfunction. Due to the 

nanoparticle- and approach-specific nature of pharmacology studies, it is beyond the scope of this 

chapter to discuss pharmacological assay specifics. Where appropriate models exist, chemotherapeutic 

efficacy can be examined in vitro. In certain cases, targeting of chemotherapeutic agents may 

be demonstrated as well, using optimized treatment/wash-out schemes in cell lines expressing the 

targeted receptor. Though nanoparticle metabolism or enzyme induction has yet to be demonstrated, 

certain nanomaterials with attractive chemistries may be subjected to phase-I/II 

metabolism and induction studies using cell-based, microsomal, and/or recombinant 

enzyme systems. 

7.3.1 SPECIAL CONSIDERATIONS 

Many of the standard methods used to evaluate biocompatibility of new molecular and chemical 

entities are fully applicable to nanoparticles. However, existing test protocols may require further 

development and laboratory validation before they become available for routine testing. Careful 

attention must be paid to potential sources of interference with analytical endpoints that may lead to 

false-positive or false-negative results. Nanoparticle interference could result from: interference 

with assay spectral measurements; inhibition/enhancement of enzymatic reactions 61,62 ; and absorption 

of reagents to nanoparticle surfaces. In the event of nanoparticle interference, additional 

sample preparation steps or alternative methods may be required. 

When evaluating the results of in vitro assays, it is important to recognize that dose–response 

relationships will not always follow a classical linear pattern. These atypical dose–response 

relationships have previously been attributed to shifts between the different mechanisms underlying 

the measured response. 63 In the case of nanomaterials, it is also important to bear in mind that 


116 

concentration-dependent changes in the physical state (e.g., aggregation state, degree of protein 

binding) may also result in apparent nonlinearity. 

Another key consideration when evaluating the results of nanoparticle research is the impact of 

dose metric (e.g., mass, particle number, surface area), sample preparation (e.g., sonication), and 

experimental conditions (e.g., exposure to light) on the interpretation of results. For example, 

surface area or particle number may be a more appropriate metric than mass when comparing 

data generated for different sized particles. This has been shown to be the case for 20- and 250-nm 

titanium dioxide nanoparticles, in which lung inflammation in rats, as assessed by percentage of 

neutrophils in lung lavage fluid, correlated with total surface area rather than mass. 64 The importance 

of experimental conditions in study design is highlighted by an investigation of functionalized 

fullerenes, demonstrating that the cytotoxicity of dendritic and malonic acid functionalized fullerenes 

to human T-lymphocytes in vitro is enhanced by photoexcitation. 41 The standardization of 

these experimental variables should limit inter-laboratory variability and make data generated 

more comparable. 

7.3.2 IN VITRO TARGET-ORGAN TOXICITY 


A recently published report from the International Life Sciences Institute Research Foundation/Risk 

Science Institute Nanomaterial Toxicity Screening Working Group 73 recommends the inclusion 

of several specific in vitro assays in a standard protocol of safety assessment. Much of the report 

focused on toxicity screening for environmental exposure to nanoparticles and thus emphasized 

environmentally relevant exposure routes. However, in addition to the in vitro examination of 

so-called portal-of-entry tissues, the report expressed the need for inclusion of potential target 

organs. The liver and kidney were selected as ideal candidates for these initial in vitro target 

organ toxicity studies, since preliminary investigations (discussed below) have identified these as 

the primary organs involved in the accumulation, processing, and eventual clearance 

of nanoparticles. 

The liver has been identified in many studies as the primary organ responsible for 

reticuloendothelial capture of nanoparticles, often due to phagocytosis by Kupffer cells. 65–67 Fluorescein 

isothiocyanate-labeled polystyrene nanoparticles and radiolabeled dendrimers, for example, 

are rapidly cleared from the systemic circulation by hepatic uptake following intravenous injection. 

4,68 Hepatic uptake has also been shown to be a primary mechanism of hepatic clearance for 

parenterally administered fullerenes, dendrimers, and quantum dots. 20,69,70 In addition to hepatic 

accumulation, nanoparticles have also been shown to have a detrimental effect on liver function ex 

vivo and alter hepatic morphology. Hepatocytes isolated from rats intravenously administered 

polyalkylcyanoacrylate nanoparticles had diminished secretion of albumin and decreased 

glucose production. 71 This alteration in albumin synthesis was also observed in freshly isolated 

rat hepatocytes exposed to polyalkylcyanoacrylate nanoparticles. Dendrimers have also been 

shown to cause liver injury. Repeated dosing of mice with polyamidoamine (PAMAM) dendrimers 

resulted in vacuolization of the hepatic parenchyma, suggesting lysosomal dysfunction. 72 

Sprague–Dawley rat hepatic primary cells and human hepatoma Hep-G2, selected for in vitro 

hepatic target organ toxicity assays, have a long history of use in toxicological evaluation. 73–75 

Hep-G2 cells were chosen since they are a readily available hepatocyte cell line with high metabolic 

activity. 76 Rat hepatic primary cells were also chosen for toxicological studies, since hepatic 

primary cells in culture are more reflective of in vivo hepatocytes with regard to enzyme expression 

and specialized functions. One survey found rat hepatic primary cells up to ten times more sensitive 

to model hepatotoxic agents than established hepatic cell lines. 77 Rat hepatic primaries represent a 

suitable alternative to human hepatic primary cells, which are scarce, costly, and suffer from 

interindividual variability. 

Preliminary pharmacokinetic studies of parenterally administered, radiolabeled carbon nanotubes, 

dendritic fullerenes, and low generation dendrimers in rodents have identified urinary 



excretion as the principal mechanism of clearance. 78–80 A variety of engineered nanoparticles, 

including actinomycin D-loaded isobutylcyanoacrylate, doxorubicin-loaded cyanoacrylate, and 

dendrimer nanoparticles, have also been shown to distribute to renal tissue following parenteral 

administration in rodents. 68,81,82 In the case of the doxorubicin-loaded cyanoacrylate nanoparticles, 

doxorubicin renal distribution was increased due to capture of the nanoparticles by glomerular 

mesangial cells. This resulted in a shift in the primary target organ from the heart to the kidney. 

Doxorubicin-induced renal injury presented as a severe proteinuria. Kidney injury has been demonstrated 

for other nanomaterials as well. Nano-zinc particles, for example, caused severe histological 

alterations in murine kidneys and Q-dots were shown to be cytotoxic to African green monkey 

kidney cells. 83,84 

The porcine renal proximal tubule cell line, LLC-PK1, was selected as a representative kidney 

cell line, since it is readily available through ATCC and has been used extensively in nephrotoxicity 

screening and mechanistic studies. 85 The SD rat hepatic primary, LLC-PK1 and Hep-G2 cells are 

adherent, which can simplify sample preparation, and can be propagated in a 96-well plate format 

suitable for high-throughput screening. The 96-well format allows for detailed concentration– 

response curves and multiple controls to be run on the same microplate. These cell lines were 

subjected to a variety of in vitro assays, described below, to evaluate cytotoxicity, mechanistic 

toxicology, and pharmacology. These assays have been selected primarily for their superior performance, 

convenience, and adaptability in evaluating this new class of biomedical agent. 

7.3.3 CYTOTOXICITY 

Cell viability of adherent cell lines can be assessed by a variety of methods. 86 These methods fall 

broadly into four categories, assays that measure: (1) loss of membrane integrity; (2) loss of 

metabolic activity; (3) loss of monolayer adherence; and (4) cell cycle analysis. Data generated 

using these various viability assays can be used to identify cell lines susceptible to nanoparticle 

toxicity and potentially give clues as to the type (i.e., cytostatic/cytotoxic) and location of cellular 

injury. Many of the cytotoxicity assays discussed below are available as commercial kits. These kits 

should be used whenever feasible since they provide an extra level of quality control. 

1. Membrane integrity assays are particularly important as a measure of cellular damage, 

since there is evidence that some cationic nanoparticles, such as amine terminated 

dendrimers, exhibit toxic effects by disrupting the cell membrane. 87 Examples of 

assays that measure membrane integrity include the trypan blue exclusion assay and 

lactate dehydrogenase (LDH) leakage assay, which measures the presence of LDH 

released into the media through cell lysis. 88,89 The LDH leakage assay was selected 

because of its sensitivity and suitability for the high-throughput, 96-well plate format. 

2. Examples of assays which measure metabolic activity include tetrazolium dye reduction, 

ATP, and 3H-thymidine incorporation assays. The 3-(4,5-dimethyl-2-thiazolyl)-2,5diphenyl-2H-tetrazolium 

bromide (MTT) reduction assay was chosen for measurement 

of metabolic activity in the assay cascade, since it does not use radioactivity, and 

historically has been proven sensitive and reliable. MTT is a yellow water-soluble 

tetrazolium dye that is metabolized by live cells to water insoluble, purple formazan 

crystals. The formazan can be dissolved in DMSO and quantified by measuring the 

absorbance of the solution at 550 nm. Comparisons between the spectra of samples 

from nanoparticle treated and untreated cells can provide a relative estimate of cytotoxicity. 

90 

The MTT assay requires a solubilization step that is not required for the newer 

generation of tetrazolium dyes that form water-soluble formazans (e.g., XTT). 

However, these analogs require an intermediate electron acceptor that is often unstable, 

adding to assay variability. Furthermore, the net negative charge of these newer analogs 


118 

limits cellular uptake, resulting in extracellular reduction. 91 MTT, with a net positive 

charge, readily crosses cell membranes and is reduced intracellularly, primarily in the 

mitochondria. Because nanoparticles have been shown to interact with cell membranes 

and could potentially interfere with the reduction of the newer generation analog via 

trans-plasma membrane electron transport, the traditional MTT assay would appear to be 

a better choice to assess cellular viability in nanoparticle cytotoxicity experiments. 

Analytes that are antioxidants, or are substrate/inhibitors of drug efflux pumps, have 

been shown to interfere with the MTT assay. 92,93 Functionalized fullerenes, which have 

not identified as efflux pump inhibitors or substrates, but do possess potent antioxidant 

activity, have been observed in our laboratory to cause MTT assay interference, resulting 

in enhanced MTT reduction and overestimation of cell viability (unpublished data). 

3. Loss of monolayer adherence to plating surfaces is often used as a marker of cytotoxicity. 

Monolayer adherence is commonly measured by staining for total protein, following 

fixation of adherent cells. This simple assay is often a very sensitive indicator of loss of 

cell viability. 55 The sulforhodamine B total protein staining assay was selected for 

determination of monolayer adherence. Advantages of this assay include the ability to 

store the fixed, stained microplates for extended periods prior to measurement, making 

the assay especially suitable for high throughput. 94 

4. Cell cycle analysis is conducted using propidium iodide staining of DNA and flow 

cytometry. 95 Flow cytometric can be used as a screening test for toxicity of chemicals. 

This method can determine the effect of nanoparticle treatment on cell cycle progression, 

as well as cell death. Cell cycle effects have been shown for a variety of nanoparticles. 

For instance, carbon nanotubes have been shown to cause G1 cell cycle arrest in human 

embryonic kidney cells, with a corresponding decrease in expression of G1-associated 

cdks and cyclins. 96 

7.3.4 OXIDATIVE STRESS 


The generation of free radicals by nanomaterials is well documented. 97,98 In most cases, the studied 

material was of ambient or industrial origin (quartz, carbon black, metal fumes, and diesel exhaust 

particles). However, engineered nanomaterials, such as fullerenes and polystyrene nanoparticles, 

have been shown to generate oxidative stress as well. 40,99,100 Lovric et al., for example, determined 

ROS to play an important role in cytotoxicity of quantum dots that have lost their protective 

coating. 101 The unique surface chemistries, large surface area, and redox active or catalytic 

contaminants (e.g., metals, quinones) of nanoparticles can facilitate ROS generation. 102 For 

example, fullerenes can perform electron transfer (phase-I pathway) or energy transfer (phase-II 

pathway) reactions with molecular oxygen following photoexcitation, 44 resulting in the formation 

of the superoxide anion radical or singlet oxygen, respectively. The superoxide anion radical can 

then undergo further reactions, such as dismutation and Fenton chemistry, to generate additional 

ROS species (e.g., % OH), resulting in cellular injury (see Scheme 7.1). 103 Evidence of fullereneinduced 

oxidative stress includes lipid peroxidation in the brains of exposed fish and treated rat liver 

microsomes. 40,104 Additional biomarkers of oxidative stress include a decrease in the reduced 

glutathione/oxidized glutathione ratio (GSH/GSSG), DNA fragmentation, and protein 

carbonyls. 105 

Biomarkers of nanoparticle-induced oxidative stress measured in our laboratory include ROS, 

lipid peroxidation products, and GSH/GSSG ratio. The fluorescent dichlorodihydroflourescein 

(DCFH) assay is used for measurement of ROS, such as hydrogen peroxide. 106 DCFH-DA is a 

ROS probe that undergoes intracellular deacetylation, followed by ROS-mediated oxidation to a 

fluorescent species, with excitation 485 nm and emission 530 nm. DCFH-DA can be used to 

measure ROS generation in the cytoplasm and cellular organelles, such as the mitochondria. The 



2GSH 

NADP + 

(1) ROS Generation 

hv 

GR 

2H 

GSSG +H2O GSHPx 

SOD 

Fullerene 

+ 

– 

O2 O2 + e 

– 

O2 – 

NADPH + H + 

(2) Loss of GSH Homeostasis 

Fe ++ 

GSH 

O 2 – 

O 2 + OH – + HO 

thiobarbituric acid reactive substances (TBARS) assay is used for measurement of lipid peroxidation 

products, such as lipid hydroperoxides and aldehydes. A molondialdehyde (MDA) 

standard curve is used for quantitation. MDA, a lipid peroxidation product, combines with thiobarbituric 

acid in a 1:2 ratio to form a fluorescent adduct, that is measured at 521 nm (excitation) 

and 552 nm (emission). TBARS are expressed as MDA equivalents. 107 The dithionitrobenzene 

(DTNB) assay is used for evaluation of glutathione homeostasis. In the DTNB assay, reduced GSH 

interacts with 5,5 0 -dithiobis(2-nitrobenzoic acid) (DTNB) to form the colored product 2-nitro-5thiobenzoic 

acid, which is measured at 415 nm, and GSSG. GSSG is then reduced by glutathione 

reductase to form reduced GSH, which is again measured by the preceding method. Pretreatment 

with thiol-masking reagent, 1-methyl-4-vinyl-pyridinium trifluoromethane sulfonate, prevents 

GSH measurement, resulting in measurement of GSSG alone. 108 

7.3.5 APOPTOSIS AND MITOCHONDRIAL DYSFUNCTION 

Nanoparticle-induced cell death can occur by either necrosis or apoptosis, processes that can be 

distinguished both morphologically and biochemically. Morphologically, apoptosis is characterized 

by perinuclear partitioning of condensed chromatin and budding of the cell membrane to 

form apoptotic bodies, whereas necrosis is characterized by cellular swelling (oncosis) and blebbing 

of the cell membrane. 109 In vitro studies have demonstrated the ability of nanoparticles, such 

as dendrimers and carbon nanotubes, to induce apoptosis. 110–112 In vitro exposure of macrophagelike 

mouse RAW 264.7 cells to cationic dendrimers led to apoptosis confirmed by morphological 

observation and the evidence of DNA cleavage. 112 Pretreatment of cells with a general caspase 

inhibitor (zVAD-fmk) reduced the apoptotic effect of the cationic dendrimer. 112 Apoptosis has also 

been observed in cultured human embryonic kidney cells (HEK293) and T lymphocytes treated 

CAT 

H 2 O 2 + O 2 

H 2 O 2 

2H 2O + O 2 

(3) Cellular Damage Lipid Peroxidation Protein/DNA Damage 

SCHEME 7.1 (1) Photoexcited fullerenes can perform electron-transfer reactions with molecular dioxygen to 

form the superoxide anion radical (O $K 

2 ). Superoxide can then undergo superoxide dismutase (SOD)-catalyzed 

dismutation to hydrogen peroxide (H 2O 2). H 2O 2 is a substrate for catalase (CAT)-and glutathione peroxidase 

(GSHPx)-catalyzed detoxification reactions. (2) The oxidation of glutathione (GSH) to form oxidized glutathione 

(GSSG) during detoxification of H2O2 can result in a loss of glutathione homeostasis. GSH can be 

regenerated by glutathione reductase (GR). (3) Alternatively, hydrogen peroxide can undergo transition metal 

(Fe CC )-catalyzed Fenton chemistry to form the highly reactive hydroxyl radical (HO % ) that is capable of 

initiating lipid peroxidation and DNA/protein oxidation. 


120 


with single walled carbon nanotubes, and in MCF-7 breast cancer cells treated with quantum 

dots. 101,110,113 

Apoptosis in mammalian cells can be initiated by four potential pathways: (1) mitochondrial 

pathway, (2) Death receptor-mediated pathway, (3) ER-mediated pathway, and (4) Granzyme 

B-mediated pathway. 114 Our laboratory has focused on caspase-3 activation in liver and kidney 

cells as a biomarker of apoptosis, since this a downstream event in all the classical apoptotic 

signaling pathways and can be measured using a fluorometric protease assay. This assay quantifies 

caspase-3 activation in vitro by measuring the cleavage of DEVD-7-amino-4-trifluoromethyl 

coumarin (AFC) to free AFC that emits a yellow-green fluorescence (l maxZ505 nm). 115 This 

initial apoptosis screen can then be followed by additional analysis, as cellular morphology 

studies using nuclear staining techniques to detect perinuclear chromatin, or agarose gel electrophoresis 

to detect DNA laddering. 116 

Evidence supports a role for ROS in generation of the mitochondrial permeability transition via 

oxidation of thiol components of the permeability transition pore complex. 117 As discussed in the 

preceding sections, nanoparticles have been shown to induce oxidative stress, and thus this ROSmediated 

pathway for induction of the mitochondrial permeability transition is a plausible apoptotic 

mechanism for nanomaterials. For instance, ambient ultrafine particulates have been shown to 

translocate to the mitochondria of RAW 264.7 murine macrophage cells, cause structural 

damage, and altered mitochondrial permeability. 98 A subsequent study demonstrated that mitochondrial 

dysfunction and apoptosis in the RAW 264.7 cells could be induced by polar compounds 

fractionated from ultrafine particles, suggesting that the mitochondrial dysfunction caused by 

ultrafines was the result of redox cycling of quinone contaminants on the surface of the particle. 118 

This link between oxidative stress, mitochondrial dysfunction, and apoptosis has also been 

observed for man-made nanoparticles. For example, metal and quantum dot engineered nanoparticles 

have both been shown to induce oxidative stress, mitochondrial dysfunction, and apoptosis in 

various in vitro models. 101,119 Water-soluble, derivatized fullerenes, which have been shown to 

accumulate in the mitochondria of HS 68 human fibroblast cells, have also been shown to induce 

apoptosis in U251 human glioma cells. 120,121 While this derivatized fullerene-induced apoptosis in 

the glioma cell line did not involve oxidative stress, mitochondrial dysfunction was not measured 

and cannot be ruled out. Mitochondrial dysfunction and apoptosis have also been observed in a 

human gastric carcinoma cell line exposed to chitosan nanoparticles. 122 Taken together, these 

observations support a role for mitochondrial dysfunction and oxidative stress in nanoparticleinduced 

apoptosis. Apart from apoptosis, mitochondrial dysfunction has long been associated 

with necrotic cell death, and represents a potential necrotic mechanism of nanoparticle-induced 

injury as well. 123 

Mitochondrial dysfunction can result from several mechanisms in addition to opening of the 

permeability transition pore complex, including uncoupling of oxidative phosphorylation, damage 

to mitochondrial DNA, disruption of the electron transport chain, and inhibition of fatty acid 

b-oxidation. 124 Methods used to detect mitochondrial dysfunction include measurement of 

ATPase activity (via luciferin–luciferase reaction), oxygen consumption (via polarographic technique), 

morphology (via electron microscopy), and membrane potential (via fluorescent probe 

analysis). 125 Our laboratory measured loss of mitochondrial membrane potential in rat hepatic 

primaries, and Hep-G2 and LLC-PK1 cell lines, using the 5,5 0 ,6,6 0 -tetrachloro-1,1 0 ,3,3 0 -tetraethylbenzimidazolcarbocyanine 

iodide (JC-1) assay, which is a convenient assay that does not require 

mitochondrial isolation or use specialized equipment. 126 This fluorescent dye partitions into the 

mitochondrial matrix as a result of the membrane potential. Concentration of JC-1 in the matrix 

results in aggregation that fluoresces at 590 nm (red). Upon loss of membrane potential, the dye 

dissipates from the matrix and can be measured, in its monomer state at emission 527 nm (green). 

The proportion of green to red fluorescence reflects the degree of mitochondrial membrane 

depolarization. 



7.3.6 PROTEOMICS AND TOXICOGENOMICS 

Proteomics and toxicogenomics are useful tools for identifying the mechanisms underlying 

toxicity. 127 Using gel electrophoresis in combination with MS identification, or gene microarray 

technology, the expression of specific pathway-responsive genes, such as Phase-II enzymes for 

oxidative stress, or cytokines for inflammation, can be identified. The delineation of these toxic 

pathways could help further refine the in vitro and in vivo study of nanomaterials, potentially 

leading to the development of novel biomarkers that could then be used in clinical and occupational 

toxicology studies. Proteomic and genomic research on biomedically relevant nanomaterials is 

presently underway, using a series of human hepatocyte, kidney, and immunological primary cells. 

7.4 IN VIVO PHARMACOKINETIC AND TOXICOLOGICAL ASSESSMENTS 

As is the case with any NCE, a thorough understanding of the properties that govern biocompatability 

is necessary to allow transition of nanomaterials to human clinical trials. Although in vitro 

toxicology studies can be informative, the obvious caveat is that phenomenon observed in vitro may 

not materialize in vivo due to differences in biological response or nanoparticle concentrations. 

Therefore, nanoparticle safety and therapeutic efficacy can only be definitively assessed by rigorous 

in vivo testing. This phase is guided in part by insights obtained from the physicochemical and 

in vitro characterization programs. 

The primary goal of in vivo studies is to evaluate nanomaterials’ pharmacokinetics, safety, and 

efficacy (see Table 7.1) in the most appropriate animal models. Preclinical toxicological and 

pharmacokinetic studies are conducted in accordance with the FDA regulatory guidance for IND 

and IDE submission. It is not within the scope of this chapter to review this regulatory guidance; 

instead, the reader is directed to the regulatory chapter of this text. While it is generally agreed that 

the current in vivo pharmacological and toxicological endpoints used for devices and small 

molecule drugs should be appropriate in assessing the safety and efficacy of biomedical nanomaterials, 

the qualities of nanomaterials that lend themselves to biomedical application, such as 

TABLE 7.1 

In Vivo Pharmacological and Toxicological Assessment of Nanomatersials 

Category Assessment 

Initial disposition study Tissue distribution 

Clearance mechanisms 

Half-life 

Systemic exposure (plasma AUC) 

Immunotoxicity 28-day screen 

Immunogenicity (repeat-dose toxicity study) 

Hypersensitivity 

Immunostimulation 

Immunosuppression 

Dose-range finding toxicity NOAEL 

STD10 

Good laboratory practices studies PK/ADME 

Expanded single dose acute toxicity 

Repeated dose toxicity 

Efficacy Targeting 

Therapeutic 

Imaging 


122 


macromolecular structure and polydispersity, could also be problematic with respect to preclinical 

characterization, as will be discussed below. Early efforts are focused on identifying and standardizing 

the analytical and toxicological methodologies that are unique to nanoparticle 

preclinical characterization. 

Several factors can influence the clinical viability of a new cancer diagnostic or therapeutic 

agent, aside from economic feasibility. These include: (1) demonstrated advantage over the current 

market standard; (2) appropriate administration route/schedule/elimination half-life; and (3) 

favorable safety profile. 

1. The in vivo characterization phase includes an assessment of nanoparticle imaging 

and/or therapeutic efficacy in animal models that most closely approximate the human 

disease state. For instance, targeting efficacy is addressed by comparing a nanoparticle 

distribution profile with that of a nontargeted nanoparticle from the same class; however, 

this approach may be prone to ambiguities due to passive targeting via the EPR effects. 

For those particles with imaging components, the signal enhancement and tissue distribution 

profile is monitored using the appropriate magnetic resonance, ultrasound, optical, 

or positron emission tomography imaging instrumentation. In all cases, the approach will 

be compared against the present market standard to provide comparable data regarding 

efficacy, pharmacokinetics, and safety. 

2. Animal studies of nanoparticles have rarely utilized the oral administration route, but 

those that have used that route have demonstrated poor intestinal absorption. For 

example, 98% of orally administered, PEG-functionalized fullerenes are eliminated in 

the feces within 48 h. Oral bioavailability of polystyrene nanoparticles were similarly 

poor, with less than 7% of 3000-, 1000-, 100-, 50-, and 3-nm sized particles absorbed. 128 

Due to this extremely low nanoparticle oral bioavailability, the majority of biomedical 

nanoparticle formulations encountered undoubtedly will be intended for parenteral 

administration. Since diagnostic and chemotherapeutic regimens are typically of short 

duration, the inconvenience of intravenous administration does not appear to be a significant 

hurdle for eventual clinical transition. 

To be a successful diagnostic or therapeutic agent, nanoparticles should be eliminated 

from the body in a reasonable timeframe. However, studies have shown that optimum 

passive targeting of tumors, by the EPR effect, also requires that nanoparticle agents 

remain in the systemic circulation for prolonged periods. 129 Therefore, there must be a 

balance between systemic residence and clearance. At present, avoidance of the RES 

system and eventual urinary clearance appears to be a formidable obstacle for many of 

the current approaches, such as iron oxide MRI contrast agents and lipidic nanoparticle 

drug delivery agents, which have been shown to undergo capture by organs of the RES 

and remain for extended periods. 130,131 The primary mechanism of nanoparticle clearance, 

as discussed previously, appears to be glomerular filtration, which is governed by 

charge, molecular weight, and degree of protein binding. 132,133 A good case study of the 

importance of molecular weight in mediating glomerular filtration, by Gillies and 

colleagues, demonstrated that dendrimer–polyethylene oxide complexes greater than 

40 kDa were cleared less readily than lower molecular weight species. Because timely 

clearance is an important drug attribute, accurate and thorough disposition studies 

are required. 

The single greatest obstacle for nanoparticle disposition studies is analytical 

methodology to quantitate nanoparticle concentrations in biological matrices, such as 

plasma and tissues. Due to their macromolecular and polydispersed nature, nanomaterials 

do not lend themselves to quantitation by traditional methods, such as HPLC and 

LC/MS, and may require alternative methods such as radiolabeling and scintillation 



counting. Since many biomedical nanoparticles will be multifunctional, and have 

imaging components, the imaging functionality of the particle could potentially be 

used for in vivo quantitation. In cases where no imaging component is present, the 

nanoparticles could be tagged with an appropriate probe to allow for imaging quantitation. 

The labeling method utilized may alter the surface properties of a 

nanoparticle, and thus affect the tissue distribution profile; comparison of alternative 

labeling methods would help identify consensus behavior. In any event, the use of 

imaging would first require validation using traditional methods such as LC/MS or 

scintillation to ensure image intensity could correlate to nanoparticle concentration in 

a linear fashion. Since many nanomaterials are electrondense, electron microscopy might 

also be used for tracking tissue distribution. 

3. The objectives of the preclinical toxicological studies are to identify target organs of 

toxicity and to aid in the selection of starting doses for phase-I human clinical trials. 

Toward this end, toxicity studies seek to determine dose ranges causing (1) no adverse 

effects (NOAEL) and (2) life-threatening toxicity (i.e., severe toxic dose 10% [STD 10]). 

Studies are performed in two mammalian species, rodent and nonrodent, with rats the 

preferred rodent species, since they exhibit the greatest concordance with human toxicities. 

134 Nanoparticle formulations are administered according to the intended clinical 

treatment cycle, with regard to schedule, duration, route, and formulation. Necropsy, 

performed on animals showing signs of morbity during the study and at study termination, 

includes comprehensive hematology, histopathology, and clinical chemistry (see 

Figure 7.3). Several preclinical studies suggest a key role for reticuloendothelial organs, 

such as liver, kidney, and bone marrow, in the uptake of nanoparticles from the systemic 

circulation. 67,69 Therefore, these organs should receive special attention with regard to 

functional and histopathological evaluation. A review of the limited in vivo safety data 

available for nanoparticles supports the scrutinizing of these tissues, as there are several 

examples of RES organ injury, including hepatotoxicity and nephrotoxicity. For 

example, intravenous administration of cationic PAMAM dendrimers at low doses has 

Brain 

Histopathology 

Pancreas 

Lymph node Esophagus 

Thyroid 

Trachea 

Pituitary 

Heart 

Thymus 

Gall bladder 

Spleen 

Lung 

Ileum 

Rectum 

Cecum 

Colon 

Lymph node Epididymis 

Prostate Seminal vesicle 

Urinary bladder Uterus 

Hardian gland Nasal sections 

Femur 

Vertebra 

Mammary gland Skin/subcuits 

BUN 

GGT 

total protein 

A/G 

Chloride 

Clinical Chemistry 

AST 

GLUC 

Albumin 

Sodium 

Salivary gland 

Parathyroid 

Adrenal 

Kidney 

Liver 

Duodenum 

Stomach 

Jejunum 

Ovary 

Testis 

Eye 

Femur 

Spinal cord 

Tongue 

ALT 

Creatinine 

Globulin 

Potassium 

Hematology 

Erthrocyte count (RBC) 

Hemoglobin (HGB) 

Hematocrit (HCT) 

Mean corpuscular volume (MCV) 

Mean corpuscular hemoglobin (MCH) 

Mean corpuscular hemoglobin conc. (MCHC) 

Platelet count (Plate) 

Reticulocyte count (RETIC) 

Total leukocyte count (WBC) 

Differential leukocyte count 

Nucleated red blood cell count 

FIGURE 7.3 Panel of tissues and chemistries to be assessed for comprehensive toxicology. 


124 

been shown to cause liver injury, as determined by histopathology and elevations in 

serum alanin aminotransferase when administered intravenously to mice; larger doses of 

the same cationic dendrimer were lethal to 100% of the mice. 135 Nephrotoxicity and 

hepatotoxicity, as determined by histopathology and serum enzyme markers, have both 

been observed in mice treated orally with nano-zinc. 72,84 

One of the potential advantages of nanoparticle drug formulations is an improved safety profile 

as a result of targeted therapy or elimination of toxic solubilization agents. For example, Baker and 

colleagues have shown that methotrexate-conjugated PAMAM dendrimers containing a folate 

receptor targeting ligand are more efficacious, and less toxic, than unformulated methotrexate 

against a murine human epithelial cancer model. 34 Abraxane is an example of a nanoparticle 

formulation of paclitaxel, presently on the market, that takes advantage of the solubilizing effect 

of albumin, eliminating the need for the toxic vehicle Cremophor EL. 136 Phase-III studies have 

shown enhanced therapeutic response of abraxane compared to Cremophor-formulated paclitaxel, 

while side effects, such as myelosuppression and peripheral neuropathy, were significantly reduced. 

As discussed above, nanoparticle drug formulations are compared against the unformulated drug to 

determine if the improvement in safety profile is realized. 

7.5 IMMUNOTOXICITY 


A growing body of evidence suggests that immunotoxicity provides a considerable contribution to 

onset and development of various disorders, including cancer and autoimmune diseases. 137–139 

Nevertheless, it was not until recently that this relatively new field of toxicology emerged as an 

important interface between the fields of novel drug design and pharmacology. Recognition of 

immunosuppressive properties of new pharmaceuticals during early drug development phase is 

very important to eliminate potentially dangerous substances from the drug pipelines. For 

example, treatment of patients diagnosed with Crohn’s disease and rheumatoid arthritis with infliximab 

and etanercept (both drugs represent neutralizing anti-TNF antibodies) resulted in increased 

incidence of tuberculosis and histoplasmosis. 140–143 Although these data did not result in withdrawal 

of any of the products from the pharmaceutical market, they helped initiate the strategy of preparing 

patients for anti-TNF therapy by screening for, and treatment of, latent tuberculosis prior to administration 

of anti-TNF medications. Immunosuppression caused by pharmaceuticals can also lead to 

the development of lymphomas and acute leukemia. 144–146 Undesirable immunostimulation caused 

by pharmacological intervention include immunogenicity, hypersensitivity, and increased risk of 

autoimmune response. The standard toxicology endpoints employed for safety assessment of new 

pharmaceuticals primarily rely on clinical chemistry and histopathological evaluation of immune 

organs and were developed several decades ago. Currently, there is an increasing demand for 

the development of new methods for immunotoxicity assessment because of drug candidates’ 

more complex structure as well as the application of new technologies in their manufacturing. 

The introduction of new molecular and immune cell biology methods into the immunotoxicology 

assessment framework is not a trivial and straightforward process. It requires not only scrupulous 

validation and standardization of the new techniques but also demonstration of the physiological 

relevance for the proposed battery of assays. These processes are expensive, time-consuming, and 

necessitate cooperation across the various pharmaceutical industry players. Unlike traditional drugs, 

multifunctional nanomaterials combine both chemistry-based and biotechnology-derived components, 

and therefore their characterization using standard methodologies requires adjustments 

and/or modification of classical experimental protocols. Below we will attempt to summarize data 

on critical aspects of immunotoxicological evaluation of nanomaterials and examine challenges in 

the application of standard methodologies for the assessments of nanoparticle safety to the 

immune system. 



7.5.1 APPLICABILITY OF STANDARD IMMUNOLOGICAL METHODS FOR NANOPARTICLE EVALUATION 

AND CHALLENGES SPECIFIC TO NANOPARTICLE CHARACTERIZATION 

Immunotoxicological evaluation of new drug candidates includes studies on both immunosuppression 

and immunostimulation, and is applicable to nanomaterials intended for use as drug candidates 

and/or drug delivery platforms. Short-term in vitro assays are being developed to allow for quick 

evaluation of nanoparticles’ biocompatibility. The in vitro immunotoxicity assay cascade includes 

the following methods: analysis of plasma protein binding by two-dimensianol polyacrylamide gel 

electrophoresis (PAGE), hemolysis, platelet aggregation, coagulation, complement activation, 

colony forming unit-granulocyte macrophage (CFU-GM), leukocyte proliferation, phagocytosis, 

cytokine secretion by macrophages, chemotaxis, oxidative burst, and evaluation of cytotoxic 

activity of NK cells. 147 In addition to these methods, our in vitro tests include sterility assessment 

based on pyrogen contamination test (L-amebocyte lysate (LAL) assay) and evaluation of microbiological 

contamination. The assay cascade is based on several regulatory documents 

recommended buy the FDA for immunotoxicological evaluation of new investigational drugs, 

medical devices, and biotechnology, derived pharmaceuticals, 148–152 as well as on ASTM and 

ISO standards developed for characterization of blood contact properties of medical 

devices. 153–155 The aim of the in vitro immunoassay cascade is to provide quick evaluation of 

nanomaterials of interest prior to initiation of more thorough in vivo studies. Challenges specific for 

immunotoxicity assessment of nanoparticulate materials are summarized below. 

7.5.1.1 Blood Contact Properties 

One important aspect of nanoparticle used for medical applications is the assurance that they will 

not cause toxicity to blood elements when injected into a patient. 

Hemolysis (i.e., damage to red blood cells) can lead to life-threatening conditions such as 

anemia, hypertension, arrhythmia, and renal failure. In our laboratory we have developed a protocol 

to evaluate hemolytic properties of nanoparticles based on the existing ASTM International standard 

used to characterize other materials. 147,156 We have identified several problems when applying 

existing protocol for nanoparticles characterization. For example, colloidal gold nanoparticles with 

size 5–50 nm have absorbance at 535 nm, which overlaps with the assay wavelength of 540 nm. 

Removal of these particles by centrifugation was required prior to sample evaluation for the 

presence of plasma-free hemoglobin to avoid false-positive results. Though it worked well for 

gold particles with size 10–50 nm, centrifugation may be problematic for other nanoparticles. 

For example, small colloidal gold particles with a size of 5 nm require higher centrifugation 

force to be removed from the supernatant. Hemoglobin has a size of 5 nm; therefore, one cannot 

exclude the possibility that ultracentrifugation of supernatant may pellet hemoglobin along with the 

gold particles and thus result in a false-negative result. Ultracentrifugation is not feasible for 

fullerenes or dendrimer particles. Analysis of polystyrene particles of 20, 50, and 80 nm revealed 

another complication. We found that particle preparation damages red blood cells; the damage is 

caused by the surfactant used during particle manufacturing, and is detected for 20- and 50-nm 

particles. For 80-nm particles, the hemolysis assay showed false-negative results due to the adsorption 

of hemoglobin by the particles. When we applied dialysis to remove surfactant, 50-nm particles 

revealed same phenomenon as 80-nm particles, i.e., they caused hemolysis, but adsorbed hemoglobin 

resulting in a false-negative result. Another potential problem with the application of 

standard hemolysis protocol for nanoparticles characterization is that metal-containing particles 

may oxidize hemoglobin and result in a change in assay OD responses. Therefore, assay procedures 

may require slight modifications depending on the particle type. In general, inclusion of particle test 

samples without blood allowed for quick assessment of the potential particle interference with the 

assay. In some instances, deduction of the result generated for blood-free particle control from that 


126 


obtained for blood-plus particle sample was possible and allowed estimation of particle potential to 

cause damage to erythrocytes. 

Blood coagulation. Blood coagulation may be affected by nanomaterials. For example, modification 

of surface chemistry has been shown to improve immunological compatibility at the 

particle–blood interface: application of polyvinyl chloride resin particles resulted in 19G4% 

decrease in platelet count, indicating platelet adhesion/aggregation and increased blood coagulation 

time; the same particle preparation coated with PEG affected neither the platelet count nor elements 

of coagulation cascade. 157 Similarly, folate-coated and PEG-coated Gd nanoparticles did not 

aggregate platelets or activate neutrophils. 158 The evaluation of nanoparticle effects on blood 

coagulation includes studies on platelet function and coagulation factors. The in vitro cascade 

includes platelet aggregation assay and four coagulation assays measuring prothrombin time, 

activated partial thromboplastin time, thrombin time, and reptilase time. 

Interaction with plasma proteins. High-resolution, 2D PAGE may be the method of choice to 

investigate plasma protein adsorption by nanoparticles. 2D PAGE has been used in several labs to 

isolate and identify plasma proteins adsorbed on the surface of stealth polycyanoacrylate particles 8 , 

liposomes 159,160 , solid lipid 161 , and iron oxide nanoparticles. 162 Proteins commonly identified on 

several types of nanomaterials include antithrombin, C3 component of complement, alpha-2macroglobulin, 

haptoglobin, plasminogen, immunoglobulins, albumin, fibrinogen, apolipoprotein, 

and transthyretin; albumin, immunoglobulins, and fibrinogen are the most abundant. Studies using 

this approach revealed that surface chemistry is important for protein adsorption. For example, 

Peracchia et al. demonstrated that coating with PEG results in approximately a fourfold reduction in 

protein binding by polycyanoacrylate particles. 163 Gessner et al. prepared polystyrene latex model 

nanoparticles with different surface charges. This study demonstrated that increasing surface charge 

density results in a quantitative increase in plasma protein adsorption, but did not show significant 

differences in the qualitative composition of the absorbed protein mixture. 164 

One of the most important step in this procedure is the separation of particles from plasma after 

incubation is complete. Ultracentrifugation was shown to be successful for isolation of iron oxide, 

solid lipid particles, and some polymer-based particles. 8,160,162 Gel filtration was applicable to 

liposomes, 159 solid lipid, and iron oxide nanoparticles. Thode and colleagues compared four 

methods for isolation of iron oxide particles, i.e., ultracentrifugation, static filtration, magnetic 

separation, and gel filtration. Depending on the method used for particle separation from bulk 

plasma, different quantities of the same proteins and different species of proteins were identified 

on the particles of the same size and surface chemistry. 162 For example, albumin was the predominant 

protein if static filtration and gel filtration were employed, while small quantities of this 

protein were found after ultracentrifugation; it was almost undetectable when magnetic separation 

was used. There was no difference in isolation of fibrinogen among the four methods. Comparable 

quantities of IgG gamma-chain were isolated using ultracentrifugation, static filtration, and 

magnetic separation, while gel filtration appeared to be inefficient in isolation of this protein. 162 

Attention has to be paid to the sample preparation to avoid artificial protein adsorption due to 

desorption during the separation, for example. Other critical steps are the number of washes to 

remove an excess of bulk plasma, the type of wash buffer, and a buffer to dislodge protein from the 

particle surface. In our lab we also found that using polypropylene low-retention tubes and pipette 

tips is crucial for isolation of particle-specific proteins (unpublished data). 

Complement activation and phagocytosis. Following intravenous administration, nanoscale 

drug carriers may suffer a drawback in that they may be taken up by cells of the mononuclear 

phagocytic system. Consequently, such uptake facilitates clearance of nanoparticulate carriers and 

associated drugs, thus leading to a decrease in drug efficacy. 165 The initial adsorption of plasma 

proteins such as components of complement and immunoglobulins promotes nanoparticles clearance. 

166,167 Therefore, the investigation of a nanoparticle’s ability to interact with and activate a 

complement and uptake by mononuclear cells seems to be one of the key assays in the preclinical 

characterization cascade. Classical immunoassays used to evaluate complement activation, such as 



the total hemolytic complement assay (CH50) and the alternative pathway (rabbit CH50 or 

APCH50), are based on the hemolysis of rabbit erythrocytes. These hemolytic assays can be 

used to measure functional activity of specific components of either pathway. The main challenge 

in applying these assays for nanomaterial characterization is the ability of nanoparticles per se to 

lyse RBCs, thus generating false-positive results. To overcome this limitation, the approach for 

evaluation of complement activation includes two techniques. One is a qualitative yes or no rapid 

screen for the presence of C3 cleavage products using western blot. The second assay is a quantitative 

evaluation of samples found positive at an initial screen for the presence of C4a, C3a, and 

C5a components of complement using a flow cytometry-based multiplex array. 

The difficulties with application of standard phagocytosis assay are: (1) light microscopy used 

in traditional phagocytosis assay 168 is not applicable to nanoparticles due to their smaller size; (2) 

when light microscopy is substituted with TEM, visualization of particle may be complicated since 

their size is similar to that of cell organelles resulting in ambiguous interpretation of TEM data, thus 

TEM is limited to electron dense metal containing particles (see Figure 7.4); (3) labeling of 

nanoparticles with fluorescent tags 9,169 is a superior approach for visualizing internalized particles, 

but should be avoided in preclinical tests as chemical attachment of fluorophore may create a new 

molecular entity with properties widely divergent from those of the original particle; (4) application 

of luminol to detect phagocytosed particles 8 provides an exceptional technique which can overcome 

all limitations listed above; however, it is not free of applicability reservations as well (e.g., 

nanoparticles may interfere with activation of luminol once it is internalized, etc.). 

CFU-GM assays. CFU-GM assays allow for the evaluation of potential nanoparticles 

interference with growth and differentiation of bone marrow stem cells into granulocyte and 

macrophages. This assay may provide valuable information for development of anti-cancer nanotechnology 

platforms. Myelosuppression is a very common dose-limiting toxicity associated with 

the use of oncology cytotoxic drugs. Incorporation of such drugs into nanoparticle carriers targeted 

to specific cancer cells may help to reduce toxicity to normal tissues, including bone marrow, and 

should be considered during initial characterization of nanocarriers. 

(a) 

2μm 

(b) 0.2μm 

FIGURE 7.4 Study of internalization of 30-nm colloidal gold nanoparticles by murine macrophage cell line 

RAW 264.7 by TEM. (a) Analysis of single cell. (b) Zoom-in analysis of the selected area in the same cell. 


128 

7.5.2 IMMUNOGENICITY 

This issue of immune stimulation by pharmaceuticals came into the forefront when biotechnologyderived 

products, especially recombinant proteins, moved toward clinical trials. Today it is evident 

that the immune system can effectively recognize biological therapeutics as foreign substances and 

build up a multi-level immune response against them. A number of factors result in the immune 

system responding to the administration of a pharmacological product, such as structure, formulation, 

folding architecture, but also degradation byproducts. 170 In addition, the route of 

administration and the dosage were shown to influence the staging and the amplitude of the 

immune system’s response. In general, immune responses to biological products could be classified 

as benign in the sense that they affect only the pharmacological efficacy of the administered 

compound. The greatest concern is the robust immune response to certain biotechnology-derived 

products that are fraught with serious clinical consequences that may even result in a fatal outcome 

due to specific recognition and elimination of the patient’s endogenous growth factors critical for 

survival. 171–173 For example, in the case of thrombopoietin, the immune response may result in the 

production of neutralizing antibodies, causing inhibition of the endogenous thrombopoietin with 

subsequent development of thrombocytopenia. 173 The patient’s immune response to recombinant 

erythropoietin product Eprex w has been reported to induce pure red cell aplasia. 171,172,174 In the latter 

example, cross-reactivity tests indicated that antibodies generated against Eprex w could also neutralize 

other forms of erythropoietin products such as Epogen w , NeoRecormon w , and Aranesp w and 

suggested that antibodies are directed against some specific conformation of the erythropoietin active 

site. 170 Although incidence of the acute pure red cell aplasia remains relatively rare, the long-term 

implications are of great concern, as over half of patients who developed the auto-antibody remained 

transfusion dependent. The potential of using multifunctional nanoparticles for medical applications 

raises a key question of whether nanoparticle materials by themselves can induce an anti-nanoparticle 

immune response, stimulate allergic reactions, or trigger synthesis of nanopreparation-specific 

IgE. One can expect that the generation of antibodies to nanoparticles will ultimately affect only 

efficacy of the particle-based product. Of greatest concern will be the immune response to particles 

functionalized with growth factors, receptors or other biological molecules, which would result in the 

formation of antibodies, neutralizing the effect of these biological molecules and leading to potential 

exclusion of both particle-linked and endogenous proteins, akin to similar effects observed with 

biotechnology-derived pharmaceuticals. There are a limited number of studies on immunogenicity of 

nanomaterials. A few of them have shown that nanoparticles may both induce nanoparticles specific 

immune response and act as adjuvants. 175–178 Although preclinical animal studies may not be 

predictive to human immune response, available data described above do suggest that the immune 

system can recognize and build an immune response against some nanoparticles. Therefore, evaluation 

of nanoparticle antigenicity is seen as an important step during preclinical development. Other 

immunogenicity characterization should include evaluation of a nanoparticles’ ability to act as an 

adjuvant and to induce allergic reactions. 



The urgency to eliminate the pain and suffering associated with cancer is fueling research into novel 

therapies at a blistering pace. Through exquisitely targeted and multifunctional approaches, nanotechnology 

in particular holds great promise for enhancing cancer therapy. This promise, however, 

will never be realized if the safety of the nanomaterial is not demonstrated to allay public concern 

and if the regulatory structure is not in place to allow proper evaluation of the science. Without such 

a framework, the return on investment will be uncertain and progress will undoubtedly be stalled. 

The effort to develop a standardized set of protocols to characterize nanomaterials and their 

biological effects will provide a foundation for future regulation and will lead to a body of knowledge 

that will guide the design of safer nanotechnology products. 



The biologic activity and toxicity of nanoscale particles are dependent on many parameters not 

typically examined for conventional small molecule therapeutics: size, shape, surface chemistry, 

stability of outer coating, agglomeration state, etc., and many conventional properties, such as 

stability or biodistribution, must be analyzed using a very different set of protocols and/or instrumentation. 

Emerging data from studies on nanoparticles engineered for medical use will build 

toward a consensus of instrumentation and experimental methodology needed to reproducibly 

determine the pharmacology and safety of these novel products. 

The greatest challenge may not be in the development of new screening technologies, but in the 

ability to promulgate an accepted set of characterization protocols throughout the Nano Bio 

industry. The Nanotechnology Characterization Laboratory at the National Cancer Institute is 

working together with the FDA, NIST, and other regulatory and standards-setting organizations 

to establish the standard assays and technologies needed for timely delivery of safe and effective 

nanotechnology products. 

REFERENCES 

1. Furumoto, K., Nagayama, S., Ogawara, K., Takakura, Y., Hashida, M., Higaki, K., and Kimura, T., 

Hepatic uptake of negatively charged particles in rats: Possible involvement of serum proteins in 

recognition by scavenger receptor, Journal of Controlled Release, 97, 133–141, 2004. 

2. Oberdorster, G., Toxicology of ultrafine particles: In vivo studies, Philosophical Transactions of the 

Royal Society of London Series A—Mathematical Physical and Engineering Sciences, 358, 

2719–2739, 2000. 

3. Ogawara, K., Yoshida, M., Higaki, K., Kimura, T., Shiraishi, K., Nishikawa, M., Takakura, Y., and 

Hashida, M., Hepatic uptake of polystyrene microspheres in rats: Effect of particle size on intrahepatic 

distribution, Journal of Controlled Release, 59, 15–22, 1999. 

4. Ogawara, K., Yoshida, M., Kubo, J., Nishikawa, M., Takakura, Y., Hashida, M., Higaki, K., and 

Kimura, T., Mechanisms of hepatic disposition of polystyrene microspheres in rats: Effects of serum 

depend on the sizes of microspheres, Journal of Controlled Release, 61, 241–250, 1999. 

5. Fenart, L., Casanova, A., Dehouck, B., Duhem, C., Slupek, S., Cecchelli, R., and Betbeder, D., 

Evaluation of effect of charge and lipid coating on ability of 60-nm nanoparticles to cross an in vitro 

model of the blood–brain barrier, Journal of Pharmacological and Experimental Therapeutics, 291, 

1017–1022, 1999. 

6. Brown, D. M., Wilson, M. R., MacNee, W., Stone, V., and Donaldson, K., Size-dependent proinflammatory 

effects of ultrafine polystyrene particles: a role for surface area and oxidative stress in the 

enhanced activity of ultrafines, Toxicology and Applied Pharmacology, 175, 191–199, 2001. 

7. Donaldson, K. and Tran, C. L., Inflammation caused by particles and fibers, Inhalation Toxicology, 

14, 5–27, 2002. 

8. Gref, R., Luck, M., Quellec, P., Marchand, M., Dellacherie, E., Harnisch, S., Blunk, T., and Muller, 

R. H., ‘Stealth’ corona-core nanoparticles surface modified by polyethylene glycol (PEG): Influences 

of the corona (PEG chain length and surface density) and of the core composition on 

phagocytic uptake and plasma protein adsorption, Colloids and Surfaces B, 18, 301–313, 2000. 

9. Leroux, J. C., Gravel, P., Balant, L., Volet, B., Anner, B. M., Allemann, E., Doelker, E., and Gurny, R., 

Internalization of poly(D,L-Lactic Acid) nanoparticles by isolated human-leukocytes and analysis of 

plasma-proteins adsorbed onto the particles, Journal of Biomedical Materials Research, 28, 471–481, 

1994. 

10. Reynolds, L. J. and Richards, R. J., Can toxicogenomics provide information on the bioreactivity of 

diesel exhaust particles? Toxicology, 165, 145–152, 2001. 


evolving from studies of ultrafine particles, Environmental Health Perspectives, 113, 823–839, 2005. 

12. Van de Coevering, R., Kreiter, R., Cardinali, F., van Koten, G., Nierengarten, J. F., and Gebbink, R., 

An octa-cationic core–shell dendrimer as a molecular template for the assembly of anionic fullerene 

derivatives, Tetrahedron Letters, 46, 3353–3356, 2005. 


130 


13. Tomalia, D. A., Naylor, A. M., and Goddard, W. A., Starburst dendrimers—molecular-level control 

of size shape, surface-chemistry, topology, and flexibility from atoms to macroscopic matter, Angewandte 

Chemie-International Edition in English, 29, 138–175, 1990. 

14. Tomalia, D. A., Baker, H., Dewald, J., Hall, M., Kallos, G., Martin, S., Roeck, J., Ryder, J., and 

Smith, P., A new class of polymers—starburst-dendritic macromolecules, Polymer Journal, 17, 

117–132, 1985. 

15. Islam, M. T., Majoros, I. J., and Baker, J. R., HPLC analysis of PAMAM dendrimer based multifunctional 

devices, Journal of Chromatography B, 822, 21–26, 2005. 

16. Binnig, G., Quate, C. F., and Gerber, C., Atomic force microscope, Physical Review Letters, 56, 

930–933, 1986. 

17. Mecke, A., Lee, D. K., Ramamoorthy, A., Orr, B. G., and Holl, M. M. B., Synthetic and natural 

polycationic polymer nanoparticles interact selectively with fluid-phase domains of DMPC lipid 

bilayers, Langmuir, 21, 8588–8590, 2005. 

18. Mecke, A., Majoros, I. J., Patri, A. K., Baker, J. R., Holl, M. M. B., and Orr, B. G., Lipid bilayer 

disruption by polycationic polymers: The roles of size and chemical functional group, Langmuir, 21, 

10348–10354, 2005. 

19. Patri, A. K., Myc, A., Beals, J., Thomas, T. P., Bander, N. H., and Baker, J. R., Synthesis and in vitro 

testing of J591 antibody–dendrimer conjugates for targeted prostate cancer therapy, Bioconjugate 

Chemistry, 15, 1174–1181, 2004. 

20. Malik, N., Wiwattanapatapee, R., Klopsch, R., Lorenz, K., Frey, H., Weener, J. W., Meijer, E. W., 

Paulus, W., and Duncan, R., Dendrimers: Relationship between structure and biocompatibility 

in vitro, and preliminary studies on the biodistribution of I-125-labelled polyamidoamine dendrimers 

in vivo, Journal of Controlled Release, 65, 133–148, 2000. 

21. Hawley, A. E., Davis, S. S., and Illum, L., Targeting of colloids to lymph nodes: Influence of 

lymphatic physiology and colloidal characteristics, Advanced Drug Delivery Reviews, 17, 

129–148, 1995. 

22. Stolnik, S., Illum, L., and Davis, S. S., Long circulating microparticulate drug carriers, Advanced 

Drug Delivery Reviews, 16, 195–214, 1995. 

23. Ishida, O., Maruyama, K., Sasaki, K., and Iwatsuru, M., Size-dependent extravasation and interstitial 

localization of polyethyleneglycol liposomes in solid tumor-bearing mice, International Journal of 

Pharmaceutics, 190, 49–56, 1999. 

24. Maeda, H., Wu, J., Sawa, T., Matsumura, Y., and Hori, K., Tumor vascular permeability and the EPR 

effect in macromolecular therapeutics: A review, Journal of Controlled Release, 65, 271–284, 2000. 

25. Brigger, I., Dubernet, C., and Couvreur, P., Nanoparticles in cancer therapy and diagnosis, Advanced 

Drug Delivery Reviews, 54, 631–651, 2002.0 

26. Kawasaki, E. S. and Player, A., Nanotechnology, nanomedicine, and the development of new, 

effective therapies for cancer, Nanomedicine: Nanotechnology, Biology and Medicine, 1, 

101–109, 2005. 

27. Torchilin, V. P., Drug targeting, European Journal of Pharmaceutical Sciences, 11, S81–S91, 2000. 

28. Prosa, T. J., Bauer, B. J., and Amis, E. J., From stars to spheres: A SAXS analysis of dilute dendrimer 

solutions, Macromolecules, 34, 4897–4906, 2001. 

29. Nisato, G., Ivkov, R., and Amis, E. J., Structure of charged dendrimer solutions as seen by smallangle 

neutron scattering, Macromolecules, 32, 5895–5900, 1999. 

30. Tsay, J. M., Doose, S., and Weiss, S., Rotational and translational diffusion of peptide-coated 

CdSe/CdS/ZnS nanorods studied by fluorescence correlation spectroscopy, Journal of the American 

Chemical Society, 128, 1639–1647, 2006. 

31. Calabretta, M., Jamison, J. A., Falkner, J. C., Liu, Y., Yuhas, B. D., Matthews, K. S., and Colvin, 

V. L., Analytical ultracentrifugation for characterizing nanocrystals and their bioconjugates, Nano 

Letters, 5, 963–967, 2005. 

32. Giddings, J. C., Field-flow fractionation: Analysis of macromolecular, colloidal, and particulate 

materials, Science, 260, 1456–1465, 1993. 

33. Lockman, P. R., Koziara, J. M., Mumper, R. J., and Allen, D. D., Nanoparticle surface charges alter 

blood–brain barrier integrity and permeability, Journal of Drug Targeting, 12, 635–641, 2004. 



34. Kukowska-Latallo, J. F., Candido, K. A., Cao, Z., Nigavekar, S. S., Majoros, I. J., Thomas, T. P., 

Balogh, L. P., Khan, M. K., and Baker, J. R., Nanoparticle targeting of anticancer drug improves 

therapeutic response in animal model of human epithelial cancer, Cancer Research, 65, 5317–5324, 

2005. 

35. Nel, A., Atmosphere: Enhanced: Air pollution-related illness: Effects of particles, Science, 308, 

804–806, 2005. 

36. Brown, D. M., Stone, V., Findlay, P., MacNee, W., and Donaldson, K., Increased inflammation and 

intracellular calcium caused by ultrafine carbon black is independent of transition metals or other 

soluble components, Occupational and Environmental Medicine, 57, 685–691, 2000. 



38. Joo, S. H., Feitz, A. J., and Waite, T. D., Oxidative degradation of the carbothioate herbicide, molinate, 

using nanoscale zero-valent iron, Environmental Science & Technology, 38, 2242–2247, 2004. 

39. Nagaveni, K., Sivalingam, G., Hegde, M. S., and Madras, G., Photocatalytic degradation of organic 

compounds over combustion-synthesized nano-TiO 2, Environmental Science & Technology, 38, 

1600–1604, 2004. 

40. Oberdorster, E., Manufactured nanomaterials (fullerenes, C60) induce oxidative stress in the brain of 

juvenile largemouth bass, Environmental Health Perspectives, 112, 1058–1062, 2004. 

41. Rancan, F., Rosan, S., Boehm, F., Cantrell, A., Brellreich, M., Schoenberger, H., Hirsch, A., and 

Moussa, F., Cytotoxicity and photocytotoxicity of a dendritic C(60) mono-adduct and a malonic acid 

C(60) tris-adduct on Jurkat cells, Journal of Photochemistry and Photobiology B, 67, 157–162, 2002. 

42. Sayes, C., Fortner, J., Guo, W., Lyon, D., Boyd, A., Ausman, K., Tao, Y., Sitharaman, B., Wilson, L., 

Hughes, J., West, J., and Colvin, V. L., The differential cytotoxicity of water-soluble fullerenes, 

Nano Letters, 4, 1881–1887, 2004. 

43. Wilson, M. R., Lightbody, J. H., Donaldson, K., Sales, J., and Stone, V., Interactions between 

ultrafine particles and transition metals in vivo and in vitro, Toxicology and Applied Pharmacology, 

184, 172–179, 2002. 

44. Yamakoshi, Y., Umezawa, N., Ryu, A., Arakane, K., Miyata, N., Goda, Y., Masumizu, T., and 

Nagano, T., Active oxygen species generated from photoexcited fullerene (C60) as potential 

medicines: O2-* versus 1O2, Journal of the American Chemical Society, 125, 12803–12809, 2003. 

45. Sonvico, F., Mornet, S., Vasseur, S., Dubernet, C., Jaillard, D., Degrouard, J., Hoebeke, J., Duguet, 

E., Colombo, P., and Couvreur, P., Folate-conjugated iron oxide nanoparticles for solid tumor 

targeting as potential specific magnetic hyperthermia mediators: synthesis, physicochemical 

characterization, and in vitro experiments, Bioconjugate Chemistry, 16, 1181–1188, 2005. 

46. Thomas, T. P., Patri, A. K., Myc, A., Myaing, M. T., Ye, J. Y., Norris, T. B., and Baker, J. R., In vitro 

targeting of synthesized antibody-conjugated dendrimer nanoparticles, Biomacromolecules, 5, 

2269–2274, 2004. 

47. Kobayashi, H. and Brechbiel, M. W., Nano-sized MRI contrast agents with dendrimer cores, 

Advanced Drug Delivery Reviews, 57, 2271–2286, 2005. 

48. Shi, X., Lesniak, W., Islam, M. T., MuNiz, M. C., Balogh, L. P., and Baker, J. J. R., Comprehensive 

characterization of surface-functionalized poly(amidoamine) dendrimers with acetamide, hydroxyl, 

and carboxyl groups, Colloids and Surfaces A, 272, 139–150, 2006. 

49. Shi, X., Majoros, I. J., Patri, A. K., Bi, X., Islam, M. T., Desai, A., Ganser, T. R., and Baker, J. R., 

Molecular heterogeneity analysis of poly(amidoamine) dendrimer-based mono- and multifunctional 

nanodevices by capillary electrophoresis, Analyst, 131, 374–381, 2006. 

50. Ferrari, M., Cancer nanotechnology: Opportunities and challenges, Nature Reviews Cancer, 5, 

161–171, 2005. 


delivery, Current Opinion in Chemical Biology, 6, 466–471, 2002. 

52. Portney, N. G. and Ozkan, M., Nano-oncology: drug delivery, imaging, and sensing, Analytical and 

Bioanalytical Chemistry, 384, 620–630, 2006. 

53. Patri, A. K., Kukowska-Latallo, J. F., and Baker, J. R., Targeted drug delivery with dendrimers: 

Comparison of the release kinetics of covalently conjugated drug and non-covalent drug inclusion 

complex, Advanced Drug Delivery Reviews, 57, 2203–2214, 2005. 


132 


54. Gillies, E. R. and Frechet, J. M., pH-responsive copolymer assemblies for controlled release of 

doxorubicin, Bioconjugate Chemistry, 16, 361–368, 2005. 

55. Brucher, E., Kinetic stabilities of gadolinium(III) chelates used as MRI contrast agents, Topics in 

Current Chemistry, 221, 103–122, 2002. 

56. Hardman, R., A toxicologic review of quantum dots: Toxicity depends on physicochemical and 

environmental factors, Environmental Health Perspectives, 114, 165–172, 2006. 



58. Dubertret, B., Skourides, P., Norris, D. J., Noireaux, V., Brivanlou, A. H., and Libchaber, A., In vivo 

imaging of quantum dots encapsulated in phospholipid micelles, Science, 298, 1759–1762, 2002. 

59. Hoshino, A., Hanaki, K., Suzuki, K., and Yamamoto, K., Applications of T-lymphoma labeled with 

fluorescent quantum dots to cell tracing markers in mouse body, Biochemical and Biophysical 

Research Communications, 314, 46–53, 2004. 

60. Voura, E. B., Jaiswal, J. K., Mattoussi, H., and Simon, S. M., Tracking metastatic tumor cell 

extravasation with quantum dot nanocrystals and fluorescence emission-scanning microscopy, 

Nature Medicine, 10, 993–998, 2004. 

61. Ueng, T. H., Kang, J. J., Wang, H. W., Cheng, Y. W., and Chiang, L. Y., Suppression of microsomal 

cytochrome P450-dependent monooxygenases and mitochondrial oxidative phosphorylation by fullerenol, 

a polyhydroxylated fullerene C60, Toxicology Letters, 93, 29–37, 1997. 

62. Shcharbin, D., Jokiel, M., Klajnert, B., and Bryszewska, M., Effect of dendrimers on pure acetylcholinesterase 

activity and structure, Bioelectrochemistry, 68, 56–59, 2006. 

63. Slikker, W., Andersen, M. E., Bogdanffy, M. S., Bus, J. S., Cohen, S. D., Conolly, R. B., David, 

R. M., Doerrer, N. G., Dorman, D. C., Gaylor, D. W., Hattis, D., Rogers, J. M., Setzer, R. W., 

Swenberg, J. A., and Wallace, K., Dose-dependent transitions in mechanisms of toxicity: Case 

studies, Toxicology and Applied Pharmacology, 201, 226–294, 2004. 

64. Oberdorster, G., Finkelstein, J. N., Johnston, C., Gelein, R., Cox, C., Baggs, R., and Elder, A. C., 

Acute pulmonary effects of ultrafine particles in rats and mice, Research Report (Health Effects 

Institute), 5–74, 2000. (disc 75–86) 

65. Bazile, D. V., Ropert, C., Huve, P., Verrecchia, T., Marlard, M., Frydman, A., Veillard, M., and 

Spenlehauer, G., Body distribution of fully biodegradable [14C]-poly(lactic acid) nanoparticles 

coated with albumin after parenteral administration to rats, Biomaterials, 13, 1093–1102, 1992. 

66. Cagle, D. W., Kennel, S. J., Mirzadeh, S., Alford, J. M., and Wilson, L. J., In vivo studies of 

fullerene-based materials using endohedral metallofullerene radiotracers, Proceedings of the 

National Academy of Sciences of the United States of America, 96, 5182–5187, 1999. 

67. Ogawara, K., Furumoto, K., Takakura, Y., Hashida, M., Higaki, K., and Kimura, T., Surface hydrophobicity 

of particles is not necessarily the most important determinant in their in vivo disposition 

after intravenous administration in rats, Journal of Controlled Release, 77, 191–198, 2001. 

68. Nigavekar, S. S., Sung, L. Y., Llanes, M., El-Jawahri, A., Lawrence, T. S., Becker, C. W., Balogh, L., 

and Khan, M. K., H-3 dendrimer nanoparticle organ/tumor distribution, Pharmaceutical Research, 

21, 476–483, 2004. 

69. Ballou, B., Lagerholm, B. C., Ernst, L. A., Bruchez, M. P., and Waggoner, A. S., Noninvasive 

imaging of quantum dots in mice, Bioconjugate Chemistry, 15, 79–86, 2004. 

70. Gharbi, N., Pressac, M., Hadchouel, M., Szwarc, H., Wilson, S. R., and Moussa, F., [60]Fullerene is a 

powerful antioxidant in vivo with no acute or subacute toxicity, Nano Letters, 5, 2578–2585, 2005. 

71. Fernandezurrusuno, R., Fattal, E., Porquet, D., Feger, J., and Couvreur, P., Evaluation of liver 

toxicological effects induced by polyalkylcyanoacrylate nanoparticles, Toxicology and Applied 

Pharmacology, 130, 272–279, 1995. 


(PAMAM) starburst dendrimers, Journal of Biomedical Materials Research, 30, 53–65, 1996. 

73. Oberdorster, G., Maynard, A., Donaldson, K., Castranova, V., Fitzpatrick, J., Ausman, K., Carter, J., 

Karn, B., Kreyling, W., Lai, D., Olin, S., Monteiro-Riviere, N., Warheit, D., and Yang, H., Principles 

for characterizing the potential human health effects from exposure to nanomaterials: elements of a 

screening strategy, Particle and Fibre Toxicology, 2, 8, 2005. 



74. Dierickx, P., Prediction of human acute toxicity by the hep G2/24-hour/total protein assay, with 

protein measurement by the CBQCA method, Alternatives to Laboratory Animals, 33, 207–213, 

2005. 

75. Knasmuller, S., Parzefall, W., Sanyal, R., Ecker, S., Schwab, C., Uhl, M., Mersch-Sundermann, V., 

Williamson, G., Hietsch, G., Langer, T., Darroudi, F., and Natarajan, A. T., Use of metabolically 

competent human hepatoma cells for the detection of mutagens and antimutagens, Mutation 

Research, 402, 185–202, 1998. 

76. Urani, C., Doldi, M., Crippa, S., and Camatini, M., Human-derived cell lines to study xenobiotic 

metabolism, Chemosphere, 37, 2785–2795, 1998. 

77. Wang, K., Shindoh, H., Inoue, T., and Horii, I., Advantages of in vitro cytotoxicity testing by using 

primary rat hepatocytes in comparison with established cell lines, Journal of Toxicological Sciences, 

27, 229–237, 2002. 

78. Gharbi, N., Pressac, M., Tomberli, V., Da Ros, T., Brettreich, M., Hadchouel, M., Arbeille, B., 

Trivin, F., Ceolin, R., Hirsch, A., Prato, M., Szwarc, H., Bensasson, R., and Moussa, F., In Fullerenes 

2000: Functionalized Fullerenes, Maggini, M., Martin, N., and Guldi, D. M., Eds., vol. 9, The 

Electrochemical Society, Pennington, NJ, 2000. 

79. Lee, C. C., MacKay, J. A., Frechet, J. M., and Szoka, F. C., Designing dendrimers for biological 

applications, Nature Biotechnology, 23, 1517–1526, 2005. 

80. Wang, H., Wang, J., Deng, X., Sun, H., Shi, Z., Gu, Z., Liu, Y., and Zhao, Y., Biodistribution of 

carbon single-wall carbon nanotubes in mice, Journal of Nanoscience and Nanotechnology, 4, 

1019–1024, 2004. 

81. Manil, L., Couvreur, P., and Mahieu, P., Acute renal toxicity of doxorubicin (adriamycin)-loaded 

cyanoacrylate nanoparticles, Pharmaceutical Research, 12, 85–87, 1995. 

82. Manil, L., Davin, J. C., Duchenne, C., Kubiak, C., Foidart, J., Couvreur, P., and Mahieu, P., Uptake 

of nanoparticles by rat glomerular mesangial cells in-vivo and in-vitro, Pharmaceutical Research, 

11, 1160–1165, 1994. 

83. Shiohara, A., Hoshino, A., Hanaki, K., Suzuki, K., and Yamamoto, K., On the cyto-toxicity caused 

by quantum dots, Microbiology and Immunology, 48, 669–675, 2004. 

84. Wang, B., Feng, W. Y., Wang, T. C., Jia, G., Wang, M., Shi, J. W., Zhang, F., Zhao, Y. L., and Chai, 

Z. F., Acute toxicity of nano- and micro-scale zinc powder in healthy adult mice, Toxicology Letters, 

161, 115–123, 2006. 

85. Toutain, H. and Morin, J. P., Renal proximal tubule cell cultures for studying drug-induced nephrotoxicity 

and modulation of phenotype expression by medium components, Renal Failure, 14, 

371–383, 1992. 

86. Mickuviene, I., Kirveliene, V., and Juodka, B., Experimental survey of non-clonogenic viability 

assays for adherent cells in vitro, Toxicology In Vitro, 18, 639–648, 2004. 

87. Hong, S., Bielinska, A. U., Mecke, A., Keszler, B., Beals, J. L., Shi, X., Balogh, L., Orr, B. G., Baker, 

J. R., and Banaszak Holl, M. M., Interaction of poly(amidoamine) dendrimers with supported lipid 

bilayers and cells: hole formation and the relation to transport, Bioconjugate Chemistry, 15, 774–782, 

2004. 

88. Decker, T. and Lohmann-Matthes, M. L., A quick and simple method for the quantitation of lactate 

dehydrogenase release in measurements of cellular cytotoxicity and tumor necrosis factor (TNF) 

activity, Journal of Immunological Methods, 115, 61–69, 1988. 

89. Korzeniewski, C. and Callewaert, D. M., An enzyme-release assay for natural cytotoxicity, Journal 

of Immunological Methods, 64, 313–320, 1983. 

90. Alley, M. C., Scudiero, D. A., Monks, A., Hursey, M. L., Czerwinski, M. J., Fine, D. L., Abbott, B. J., 

Mayo, J. G., Shoemaker, R. H., and Boyd, M. R., Feasibility of drug screening with panels of human 

tumor cell lines using a microculture tetrazolium assay, Cancer Research, 48, 589–601, 1988. 

91. Berridge, M. V., Herst, P. M., and Tan, A. S., Tetrazolium dyes as tools in cell biology: New insights 

into their cellular reduction, Biotechnology Annual Review, 11, 127–152, 2005. 

92. Natarajan, M., Mohan, S., Martinez, B. R., Meltz, M. L., and Herman, T. S., Antioxidant compounds 

interfere with the 3, Cancer Detection and Prevention, 24, 405–414, 2000. 

93. Vellonen, K. S., Honkakoski, P., and Urtti, A., Substrates and inhibitors of efflux proteins interfere 

with the MTT assay in cells and may lead to underestimation of drug toxicity, European Journal of 

Pharmaceutical Sciences, 23, 181–188, 2004. 


134 


94. Voigt, W., Sulforhodamine B assay and chemosensitivity, Methods in Molecular Medicine, 110, 

39–48, 2005. 

95. Tuschl, H. and Schwab, C. E., Flow cytometric methods used as screening tests for basal toxicity of 

chemicals, Toxicology In Vitro, 18, 483–491, 2004. 

96. Cui, D., Tian, F., Ozkan, C. S., Wang, M., and Gao, H., Effect of single wall carbon nanotubes on 

human HEK293 cells, Toxicology Letters, 155, 73–85, 2005. 

97. Tao, F., Gonzalez-Flecha, B., and Kobzik, L., Reactive oxygen species in pulmonary inflammation 

by ambient particulates, Free Radical Biology and Medicine, 35, 327–340, 2003. 

98. Li, N., Sioutas, C., Cho, A., Schmitz, D., Misra, C., Sempf, J., Wang, M., Oberley, T., Froines, J., and 

Nel, A., Ultrafine particulate pollutants induce oxidative stress and mitochondrial damage, Environmental 

Health Perspectives, 111, 455–460, 2003. 

99. Fernandez-Urrusuno, R., Fattal, E., Rodrigues, J. M., Feger, J., Bedossa, P., and Couvreur, P., Effect 

of polymeric nanoparticle administration on the clearance activity of the mononuclear phagocyte 

system in mice, Journal of Biomedical Materials Research, 31, 401–408, 1996. 

100. Fernandez-Urrusuno, R., Fattal, E., Feger, J., Couvreur, P., and Therond, P., Evaluation of hepatic 

antioxidant systems after intravenous administration of polymeric nanoparticles, Biomaterials, 18, 

511–517, 1997. 

101. Lovric, J., Cho, S. J., Winnik, F. M., and Maysinger, D., Unmodified cadmium telluride quantum 

dots induce reactive oxygen species formation leading to multiple organelle damage and cell death, 

Chemical Biology, 12, 1227–1234, 2005. 

102. Risom, L., Moller, P., and Loft, S., Oxidative stress-induced DNA damage by particulate air pollution, 

Mutation Research, 592, 119–137, 2005. 

103. Grisham, M. B. and McCord, J. M., Chemistry and cytotoxicity of reactive oxygen metabolites, In 

Physiology of Oxygen Radicals, Taylor, A. E., Matalon, S., and Ward, P., Eds., American Physiology 

Society, Bethesda, MD, pp. 1–18, 1986. 

104. Kamat, J. P., Devasagayam, T. P. A., Priyadarsini, K. I., and Mohan, H., Reactive oxygen species 

mediated membrane damage induced by fullerene derivatives and its possible biological implications, 

Toxicology, 155, 55–61, 2000. 

105. Dotan, Y., Lichtenberg, D., and Pinchuk, I., Lipid peroxidation cannot be used as a universal 

criterion of oxidative stress, Progress in Lipid Research, 43, 200–227, 2004. 

106. Black, M. J. and Brandt, R. B., Spectrofluorometric analysis of hydrogen peroxide, Analytical 

Biochemistry, 58, 246–254, 1974. 

107. Dubuisson, M. L., de Wergifosse, B., Trouet, A., Baguet, F., Marchand-Brynaert, J., and Rees, J. F., 

Antioxidative properties of natural coelenterazine and synthetic methyl coelenterazine in rat hepatocytes 

subjected to tert-butyl hydroperoxide-induced oxidative stress, Biochemical Pharmacology, 

60, 471–478, 2000. 

108. Shaik, I. H. and Mehvar, R., Rapid determination of reduced and oxidized glutathione levels using a 

new thiol-masking reagent and the enzymatic recycling method: application to the rat liver and bile 

samples, Analytical and Bioanalytical Chemistry, 385 (1), 105–113, 2006. 

109. Van Cruchten, S. and Van den Broeck, W., Morphological and biochemical aspects of apoptosis, 

oncosis and necrosis, Anatomia, Histologia, Embryologia: Journal of Veterinary Medicine Series C, 

31, 214–223, 2002. 

110. Bottini, M., Bruckner, S., Nika, K., Bottini, N., Bellucci, S., Magrini, A., Bergamaschi, A., and 

Mustelin, T., Multi-walled carbon nanotubes induce T lymphocyte apoptosis, Toxicology Letters, 

160, 121–126, 2006. 



112. Kuo, J. H., Jan, M. S., and Chiu, H. W., Mechanism of cell death induced by cationic dendrimers in RAW 

264.7 murine macrophage-like cells, Journal of Pharmacy and Pharmacology, 57, 489–495, 2005. 



114. Wang, Z. B., Liu, Y. Q., and Cui, Y. F., Pathways to caspase activation, Cell Biology International, 

29, 489–496, 2005. 

115. Gurtu, V., Kain, S. R., and Zhang, G., Fluorometric and colorimetric detection of caspase activity 

associated with apoptosis, Analytical Biochemistry, 251, 98–102, 1997. 



116. Loo, D. T. and Rillema, J. R., Measurement of cell death, Methods in Cell Biology, 57, 251–264, 1998. 

117. Le Bras, M., Clement, M. V., Pervaiz, S., and Brenner, C., Reactive oxygen species and the mitochondrial 

signaling pathway of cell death, Histology and Histopathology, 20, 205–219, 2005. 

118. Xia, T., Korge, P., Weiss, J. N., Li, N., Venkatesen, M. I., Sioutas, C., and Nel, A., Quinones and 

aromatic chemical compounds in particulate matter induce mitochondrial dysfunction: implications 

for ultrafine particle toxicity, Environmental Health Perspectives, 112, 1347–1358, 2004. 

119. Hussain, S. M., Hess, K. L., Gearhart, J. M., Geiss, K. T., and Schlager, J. J., In vitro toxicity of 

nanoparticles in BRL 3A rat liver cells, Toxicology In Vitro, 19, 975–983, 2005. 

120. Foley, S., Crowley, C., Smaihi, M., Bonfils, C., Erlanger, B. F., Seta, P., and Larroque, C., Cellular 

localisation of a water-soluble fullerene derivative, Biochemical and Biophysical Research Communications, 

294, 116–119, 2002. 

121. Isakovic, A., Markovic, Z., Todorovic-Markovic, B., Nikolic, N., Vranjes-Djuric, S., Mirkovic, M., 

Dramicanin, M., Harhaji, L., Raicevic, N., Nikolic, Z., and Trajkovic, V., Distinct cytotoxic 

mechanisms of pristine versus hydroxylated fullerene, Toxicological Sciences, 91 (1), 173–183, 2006. 

122. Qi, L. F., Xu, Z. R., Li, Y., Jiang, X., and Han, X. Y., In vitro effects of chitosan nanoparticles on 

proliferation of human gastric carcinoma cell line MGC803 cells, World Journal of Gastroenterology, 

11, 5136–5141, 2005. 

123. Hirsch, T., Susin, S. A., Marzo, I., Marchetti, P., Zamzami, N., and Kroemer, G., Mitochondrial 

permeability transition in apoptosis and necrosis, Cell Biology and Toxicology, 14, 141–145, 1998. 

124. Amacher, D. E., Drug-associated mitochondrial toxicity and its detection, Current Medicinal Chemistry, 

12, 1829–1839, 2005. 

125. Gogvadze, V., Orrenius, S., and Zhivotovsky, B., Analysis of mitochondrial dysfunction during cell 

death. In Current Protocols in Toxicology, Wiley, New York, pp. 10.1–2.10.27, 2004. 

126. Guo, W.-X., Pye, Q. N., Williamson, K. S., Stewart, C. A., Hensley, K. L., Kotake, Y., Floyd, R. A., 

and Broyles, R. H., Mitochondrial dysfunction in choline deficiency-induced apoptosis in cultured 

rat hepatocytes, Free Radical Biology and Medicine, 39, 641–650, 2005. 

127. Merrick, B. A. and Madenspacher, J. H., Complementary gene and protein expression studies and 

integrative approaches in toxicogenomics, Toxicology and Applied Pharmacology, 207, 189–194, 2005. 

128. Jani, P., Halbert, G. W., Langridge, J., and Florence, A. T., Nanoparticle uptake by the rat gastrointestinal 

mucosa: Quantitation and particle size dependency, Journal of Pharmacy and Pharmacology, 42, 

821–826, 1990. 

129. Duncan, R., Polymer conjugates for tumour targeting and intracytoplasmic delivery. The EPR effect 

as a common gateway? Pharmaceutical Science & Technology Today, 2, 441–449, 1999. 

130. Briley-Saebo, K., Hustvedt, S. O., Haldorsen, A., and Bjornerud, A., Long-term imaging effects in 

rat liver after a single injection of an iron oxide nanoparticle based MR contrast agent, Journal of 

Magnetic Resonance Imaging, 20, 622–631, 2004. 

131. Cahouet, A., Denizot, B., Hindre, F., Passirani, C., Heurtault, B., Moreau, M., Le Jeune, J., and 

Benoit, J., Biodistribution of dual radiolabeled lipidic nanocapsules in the rat using scintigraphy and 

gamma counting, International Journal of Pharmaceutics, 242, 367–371, 2002. 

132. Deen, W. M., Lazzara, M. J., and Myers, B. D., Structural determinants of glomerular permeability, 

American Journal of Physiology—Renal Physiology, 281, F579–F596, 2001. 

133. Walton, K., Dorne, J. L., and Renwick, A. G., Species-specific uncertainty factors for compounds 

eliminated principally by renal excretion in humans, Food and Chemical Toxicology, 42, 261–274, 2004. 

134. Olson, H., Betton, G., Robinson, D., Thomas, K., Monro, A., Kolaja, G., Lilly, P., Sanders, J., Sipes, 

G., Bracken, W., Dorato, M., Van Deun, K., Smith, P., Berger, B., and Heller, A., Concordance of the 

toxicity of pharmaceuticals in humans and in animals, Regulatory Toxicology and Pharmacology, 

32, 56–67, 2000. 

135. Neerman, M. F., Zhang, W., Parrish, A. R., and Simanek, E. E., In vitro and in vivo evaluation of a 

melamine dendrimer as a vehicle for drug delivery, International Journal of Pharmaceutics, 281, 

129–132, 2004. 

136. Harries, M., Ellis, P., and Harper, P., Nanoparticle albumin-bound paclitaxel for metastatic breast 

cancer, Journal of Clinical Oncology, 23, 7768–7771, 2005. 

137. Descotes, J., Importance of immunotoxicity in safety assessment: a medical toxicologist’s perspective, 

Toxicology Letters, 149, 103–108, 2004. 


136 


138. Dobrovolskaia, M. A. and Kozlov, S. V., Inflammation and cancer: When NF-kappa B amalgamates 

the perilous partnership, Current Cancer Drug Targets, 5, 325–344, 2005. 

139. Merk, H. F., Sachs, B., and Baron, J., The skin: Target organ in immunotoxicology of smallmolecular-weight 

compounds, Skin Pharmacology and Applied Skin Physiology, 14, 419–430, 2001. 

140. Myers, A., Clark, J., and Foster, H., Tuberculosis and treatment with infliximab, New England 

Journal of Medicine, 346, 625–626, 2002. 

141. Keane, J., Gershon, S., Wise, R. P., Mirabile-Levens, E., Kasznica, J., Schwieterman, W. D., Siegel, 

J. N., and Braun, M. M., Tuberculosis associated with infliximab, a tumor necrosis factor (alpha)neutralizing 

agent, New England Journal of Medicine, 345, 1098–1104, 2001. 

142. Zhang, Z., Correa, H., and Begue, R. E., Tuberculosis and treatment with infliximab, New England 

Journal of Medicine, 346, 623–626, 2002. 

143. Lim, W. S., Powell, R. J., and Johnston, I. D., Tuberculosis and treatment with infliximab, New 

England Journal Medicine, 346, 623–626, 2002. 

144. Nart, D., Nalbantgil, S., Yagdi, T., Yilmaz, F., Hekimgil, M., Yuce, G., and Hamulu, A., Primary 

cardiac lymphoma in a heart transplant recipient, Transplant Proceedings, 37, 1362–1364, 2005. 

145. Caillard, S., Pencreach, E., Braun, L., Marcellin, L., Jaegle, M. L., Wolf, P., Parissiadis, A., 

Hannedouche, T., Gaub, M. P., and Moulin, B., Simultaneous development of lymphoma in recipients 

of renal transplants from a single donor: Donor origin confirmed by human leukocyte antigen 

staining and microsatellite analysis, Transplantation, 79, 79–84, 2005. 

146. Karakus, S., Ozyilkan, O., Akcali, Z., Demirhan, B., and Haberal, M., Acute myeloid leukemia 4 

years after Kaposi’s sarcoma in a renal transplant recipient, Onkologie, 27, 163–165, 2004. 

147. http://ncl.cancer.gov/working_assay-cascade.asp 

148. FDA/CDER, Guidance for industry. Immunotoxicology evaluation of investigational new drugs, 2002, 

http://www.fda.gov/Cder/guidance/. 

149. FDA/CDER, ICH S8: Immunotoxicity studies for human pharmaceuticals (draft), 2004, http://www. 

fda.gov/Cder/guidance/. 

150. FDA/CBER/CDER, Guidance for industry. Developing medical imaging drug and biological 

products. Part 1: Conducting safety assessments, 2004, http://www.fda.gov/Cder/guidance/. 

151. FDA/CDER/CBER, Guidance for industry. ICH S6: Preclinical safety evaluation of biotechnologyderived 

pharmaceuticals, 1997, http://www.fda.gov/Cder/guidance/. 

152. FDA/CDRH, Guidance for industry and FDA reviewers. Immunotoxicity testing guidance, May 

1999, http://www.fda.gov/cdrh/guidance.html. 

153. ANSI/AAMI/ISO, 10993-4: Biological evaluation of medical devices. Part 4: Selection of tests for 

interaction with blood, 2002, http://www.iso.org/iso/en/CatalogueListPage.CatalogueList. 

154. ASTM, Standard practice. F1906-98: Evaluation of immune responses in biocompatibility testing 

using ELISA tests, lymphocyte proliferation, and cell migration, 2003, http://www.astm.org/cgibin/ 

SoftCart.exe/NEWSITE_JAVASCRIPT/index.shtml?L+mystore+zdrn2970+1157995562. 

155. ASTM, Standard practice. F748-98: Selecting generic biological test methods for 

materials and devices, http://www.astm.org/cgibin/SoftCart.exe/NEWSITE_JAVASCRIPT/index. 

shtml?L+mystore+zdrn2970+1157995562. 

156. ASTM, F756-00: Standard practice for assessment of hemolytic properties of materials, 2000, 

http://www.astm.org/cgibin/SoftCart.exe/NEWSITE_JAVASCRIPT/index.shtml?L+mystore+zdrn 

2970+1157995562. 

157. Balakrishnan, B., Kumar, D. S., Yoshida, Y., and Jayakrishnan, A., Chemical modification of 

poly(vinyl chloride) resin using poly(ethylene glycol) to improve blood compatibility, Biomaterials, 

26, 3495–3502, 2005. 

158. Oyewumi, M. O., Yokel, R. A., Jay, M., Coakley, T., and Mumper, R. J., Comparison of cell uptake, 

biodistribution and tumor retention of folate-coated and PEG-coated gadolinium nanoparticles in 

tumor-bearing mice, Journal of Controlled Release, 95, 613–626, 2004. 

159. Diederichs, J. E., Plasma protein adsorption patterns on liposomes: establishment of analytical 

procedure, Electrophoresis, 17, 607–611, 1996. 

160. Harnisch, S. and Muller, R. H., Plasma protein adsorption patterns on emulsions for parenteral 

administration: establishment of a protocol for two-dimensional polyacrylamide electrophoresis, 

Electrophoresis, 19, 349–354, 1998. 



161. Goppert, T. M. and Muller, R. H., Alternative sample preparation prior to two-dimensional electrophoresis 

protein analysis on solid lipid nanoparticles, Electrophoresis, 25, 134–140, 2004. 

162. Thode, K., Luck, M., Semmler, W., Muller, R. H., and Kresse, M., Determination of plasma protein 

adsorption on magnetic iron oxides: sample preparation, Pharmaceutial Research, 14, 905–910, 1997. 

163. Peracchia, M. T., Harnisch, S., Pinto-Alphandary, H., Gulik, A., Dedieu, J. C., Desmaele, D., 

d’Angelo, J., Muller, R. H., and Couvreur, P., Visualization of in vitro protein-rejecting properties 

of PEGylated stealth (R) polycyanoacrylate nanoparticles, Biomaterials, 20, 1269–1275, 1999. 

164. Gessner, A., Lieske, A., Paulke, B. R., and Muller, R. H., Influence of surface charge density on 

protein adsorption on polymeric nanoparticles: Analysis by two-dimensional electrophoresis, 

European Journal of Pharmaceutics and Biopharmaceutics, 54, 165–170, 2002. 



166. Blunk, T., Hochstrasser, D. F., Sanchez, J. C., Muller, B. W., and Muller, R. H., Colloidal carriers for 

intravenous drug targeting: Plasma protein adsorption patterns on surface-modified latex particles 

evaluated by two-dimensional polyacrylamide gel electrophoresis, Electrophoresis, 14, 1382–1387, 

1993. 

167. Diluzio, N. R. and Zilversmit, D. B., Influence of exogenous proteins on blood clearance and tissue 

distribution of colloidal gold, American Journal of Physiology, 180, 563–565, 1955. 

168. Campbell, P. A., Canono, B. P., and Drevets, D. A., Measurement of bacterial ingestion and killing 

by macrophages, Current Protocols in Immunology, Suppl. 12, 14.6.1–14.6.13, 1994. 

169. Prabha, S., Zhou, W. Z., Panyam, J., and Labhasetwar, V., Size-dependency of nanoparticlemediated 

gene transfection: studies with fractionated nanoparticles, International Journal of Pharmaceutics, 

244, 105–115, 2002. 

170. Chamberlain, P. and Mire-Sluis, A. R., An overview of scientific and regulatory issues for the 

immunogenicity of biological products, Developments in Biological Standardization (Basel), 112, 

3–11, 2003. 

171. Casadevall, N., Antibodies against rHuEPO: native and recombinant, Nephrology Dialysis Transplantation, 

17(Suppl. 5), 42–47, 2002. 

172. Casadevall, N., Nataf, J., Viron, B., Kolta, A., Kiladjian, J. J., Martin-Dupont, P., Michaud, P., Papo, 

T., Ugo, V., Teyssandier, I., Varet, B., and Mayeux, P., Pure red-cell aplasia and antierythropoietin 

antibodies in patients treated with recombinant erythropoietin, New England Journal of Medicine, 

346, 469–475, 2002. 

173. Koren, E., Zuckerman, L. A., and Mire-Sluis, A. R., Immune responses to therapeutic proteins in 

humans—clinical significance, assessment and prediction, Current Pharmaceutical Biotechnology,3, 

349–360, 2002. 

174. Swanson, S. J., Ferbas, J., Mayeux, P., and Casadevall, N., Evaluation of methods to detect and 

characterize antibodies against recombinant human erythropoietin, Nephron Clinical Practice, 96, 

c88–c95, 2004. 

175. Andreev, S. M., Babakhin, A. A., Petrukhina, A. O., Romanova, V. S., Parnes, Z. N., and Petrov, 

R. V., Immunogenetic and allergenic activities conjugates of fullere with amino acids and protein, 

Doklady Akademii Nauk, 370, 261–264, 2000. 

176. Braden, B. C., Goldbaum, F. A., Chen, B. X., Kirschner, A. N., Wilson, S. R., and Erlanger, B. F., Xray 

crystal structure of an anti-Buckminsterfullerene antibody Fab fragment: Biomolecular recognition 

of C-60, Proceedings of the National Academy of Sciences of the United States of America, 97, 

12193–12197, 2000. 

177. Chen, B. X., Wilson, S. R., Das, M., Coughlin, D. J., and Erlanger, B. F., Antigenicity of fullerenes: 

antibodies specific for fullerenes and their characteristics, Proceedings of the National Academy of 

Sciences of the United States of America, 95, 10809–10813, 1998. 

178. Masalova, O. V., Shepelev, A. V., Atanadze, S. N., Parnes, Z. N., Romanova, V. S., Vol’pina, O. M., 

Semiletov Iu, A., and Kushch, A. A., Immunostimulating effect of water-soluble fullerene derivatives—perspective 

adjuvants for a new generation of vaccine, Doklady Akademii Nauk, 369, 

411–413, 1999. 


8 

CONTENTS 

Nanotechnology: Regulatory 

Perspective for Drug 

Development in Cancer 

Therapeutics 

N. Sadrieh and T. J. Miller 

8.1 Introduction ......................................................................................................................... 139 

8.2 FDA Experience with Nanotechnology Product Applications........................................... 140 

8.3 FDA Definition of Nanotechnology ................................................................................... 143 

8.4 FDA Initiatives in Nanotechnology .................................................................................... 143 

8.5 Research at the FDA on Nanotechnology .......................................................................... 145 

8.6 Regulatory Considerations for Nanotechnology Products ................................................. 145 

8.6.1 Jurisdiction of Nanotechnology Products ............................................................... 145 

8.6.2 Existing FDA Regulations That Apply to Nanomaterial-Containing Products..... 146 

8.7 Scientific Considerations for the Development of Nanotechnology Products .................. 147 

8.7.1 Characterization of Nanoparticles........................................................................... 147 

8.7.2 Safety Considerations.............................................................................................. 149 

8.8 Conclusions ......................................................................................................................... 149 

8.9 Trademark Listing ............................................................................................................... 151 

References .................................................................................................................................... 151 


In the last decade, remarkable scientific progress in the development of novel, nanoscale technologies 

has generated rising expectations for this enabling technology to advance our efforts in the 

prevention, diagnosis and treatment of human disease. These nanotechnologies, commonly referred 

to as nanomedicines when applied to the treatment, diagnosis, monitoring, and control of biological 

systems, incorporate a multitude of diverse and often unique nanosized products including new 

and/or improved therapeutic drugs, diagnostic/surgical devices, and targeted drug delivery 

systems. 1 The list of nanoparticle medicines approved and/or currently being developed is quite 

large; these products range from early forms of drug-encapsulating liposomes (e.g., liposomes 

containing doxorubicin for treatment of breast and ovarian cancer) and simple polymeric structures 

(e.g., actinomycin D adsorbed on poly(methylcryanoacrylate) (PMCA) and poly(ethylcryanoacrylate) 

(PECA) nanosized particles) 2 to more intricate multifunctional “nanoplatforms” and 


139

140 


diagnostic devices, including polymeric dendrimers, nanocrystal quantum dots, carbon-based fullerenes, 

and nanoshells, to name a few (see Table 8.1). The vast majority of these products are being 

developed to improve early detection of cancer and/or to provide alternative clinical therapeutic 

approaches to fighting cancer, some of these products have been approved by the Food and Drug 

Administration for clinical use. 75 

Although there have been many articles published on the in vitro toxicity of certain nanomaterials, 

very little in vivo safety data evaluating the general toxicity of these novel products is 

available in the published literature. Many in vitro assays (cell-based, organelle-based, and cell/organelle-free 

systems) have been used to evaluate the cytotoxicity of nanomaterials to help predict 

the relative safety and general biocompatibility of the nanoproducts in vivo. Several such studies 

include viability assays in vascular endothelial cells exposed to C 60; 76 fullerene induction/inhibition 

of oxidative stress and lipid peroxidation in isolated microsomes, dermal fibroblasts, and 

cultured cortical neurons; 77–79 cytotoxicity of metal oxide nanoparticles to BRL 3A liver cells; 80 

and macrophage apoptosis after exposure to carbon nanomaterials and ultrafine particles. 81–83 

Although a vast amount of useful information has been obtained from these various in vitro 

models, the relevance and application of these findings to human drug safety remain uncertain 

until validated with in vivo studies with identical nanomaterials. 

Targeted in vivo efficacy data is readily available for many of the novel, potential nanotherapeutics; 

however, necessary in vivo data indicating mechanism of action, general 

pharmacology/toxicity profiles, and optimal product characterization efforts rarely appear in the 

published literature. Much of our current understanding of the biocompatibility of nanomaterials 

comes from focused inhalation and dermal studies using nanosized products and ultrafine contaminants, 

as a model of environmental and occupational risk. 84–87 However, there are questions 

regarding the level of applicability of these studies to human drug safety assessment due to the 

variability in size and purity of the chemical or structural composition reported or not reported in 

these studies, the route of primary exposure, and the complex differences in the structure and 

function of the affected organs. 88 In addition, other in vivo studies, which identify intrinsic antioxidant 

and oxidative stress-inducing properties of unsubstituted fullerenes in a rodent and fish 

species, respectively, have important but limited value in drug discovery in the absence of systemic, 

whole animal toxicology data 67,89 . 

While the enthusiasm and expectations for success in utilizing nanomaterials and nanometersized 

structures in biomedical applications grow, so do the concerns regarding the potential impact 

of such products on the environment and human health. 88 The absence of available safety information 

for a vast majority of these novel nanosized materials and nanoparticlized drugs in 

biological systems only further supports the need for additional research into any potential toxic 

mechanisms in humans. The questions that appear most often in deciding the safety of any nanoscale 

material include: 

† Are nanomaterials and nanotechnology new to the United States Food and Drug 

Administration (FDA)? 

† How do existing regulations ensure the development of safe and effective nanotechnology-based 

drugs and how is the FDA preparing to deal with nanotechnology? 

† What are the scientific considerations that raise unique concerns for nanotechnologybased 

therapeutics? 

8.2 FDA EXPERIENCE WITH NANOTECHNOLOGY PRODUCT APPLICATIONS 

Historically, the FDA has encountered and approved many products with particulate materials 

characteristic of a nanomedicine or nanoproduct (e.g., small size, mechanism of delivery, increased 


TABLE 8.1 

Nanomedicines: Drugs/Devices FDA Approved and in Development 

Name Structure Company Size (nm) Phase Indication Reference a,b 

Magnevist wc 

PIO Shering AG !1 APP MRI tumor imaging 3–5 

Feridex wc 

SPIO Adv.Magnetics 120–180 APP MRI tumor imaging 5–9 

Rapamune wc 

Nanocrystal drug Wyeth 100–1,000 APP Immunosuppressant agent 10–12 

Emend wc 

Nanocrystal drug Merck 100–1,000 APP Nausea 13,14 

Doxil wc 

Liposome-drug encapsulation Ortho Biotech z100 APP Metastatic ovarian cancer 15–17 

Abraxane wc 

Nanoparticulate albumin American Pharmaceutical z130 APP Non-small-cell lung cancer, breast cancer, 18–20 

Partners 

others 

SilvaGard w 

Silver nanoparticles AcryMed z10 APP Antimicrobial surface treatment (medical 

devices) 

21–23 

AmBisome wc 

Liposome-drug encapsulation Gilead Science 45–80 APP Fungal infections 24–26 

Leunesse w 

Solid lipid nanoparticles Aano-therapeutics N/A On market Cosmetics N/A 

NB-001 d 

Nanoemulsion NanoBio Corp z150 Phase II Topical microbicide for herpes labialis 27–30 

VivaGel w 

Dendrimer Starpharma Holdings Ltd. N/A Phase I Vaginal microbicide (STI and HIV 

prevention) 

31–34 

INGN-401 c 

Drug–liposome complex Introgen Therapeutics N/A Phase I Metastatic lung cancer 35–38 

Combidex wc 

USPIO Advanced Magnetics 20–50 NDA filed MRI tumor imaging 39–42 

IT-101 c 

Drug–polymer complex Insert Therapeutics N/A IND filed Metastatic solid tumors 43,44 

Dendrimers c 

Polymeric spheres, multifunctional Dendritech Nanotech., O1 to!500 IDV Diagnostic, contrast agents, microbicide, 45–54 

Starpharma 

drug delivery, B neutron capture 

Nanoshells c 

Metal shell, SiO2 core Nanospectra 100–200 IDV Photo-thermal ablation therapy /imaging 

tumors 

55–58 

Fullerenes c 

Carbon sphere, C60, multifunctional C Sixty 1 IDV MRI contrast agent, antiviral, antioxidant 59–67 

Quantum dots c 

CdSe a nanocrystal Evident Technologies 2–10 IDV Optical imaging, tumor imaging, 

photosensitizer 

68–74 

APP, approved; IDV, in development; PIO, paramagnetic iron oxide; SPIO, superparamagnetic iron oxide; USPIO, ultrasmall superparamagnetic iron oxide; STI, sexually transmitted 

infection; HIV, human immunodeficiency virus; N/A, not available. 

a 

Most typical semiconductor metal used, however can also be composed of nearly any other semiconductor metal (e.g., CdS, CdTe, ZnS, PbS, etc.). 

b 

References do not necessarily refer to the product described, and may reflect general properties about the product. 

c 

Drugs with proposed or proven diagnostic or therapeutic benefit to cancer patients. 

d 

Unable to locate peer-reviewed publications. References listed are press releases found during internet search. 


Regulatory Perspective in Nanotechnology Products 141

142 


surface area, specialized function related to size and increased surface area, etc.). Although the field 

of nanomedicine may be new, the submission of applications for products containing novel dosage 

forms or drug delivery systems, including nanomaterials, nanoparticlized drugs and medical 

devices, is not particularly new to this agency. Products are reviewed on a product-by-product 

basis, and as such, various concepts of risk management (i.e., risk identification, risk analysis, risk 

control, etc.) are often implemented to support the drug review process. 

The FDA is aware of several FDA regulated products that employ nanotechnology. However, 

to date, few manufacturers of regulated products have claimed the use of nanotechnology in the 

manufacture of their products or made any nanotechnology claims for the finished product. 

The FDA is aware that a few cosmetic products and sunscreens claim to contain nanoparticles 

to increase the product stability, modify release of ingredients, or change the opacity of the 

product by using nanoparticles of titanium dioxide and zinc oxide. Similarly, several drugs 

such as Gd-DTPA (Magnevist w , tumor contrast agent), 3–5 ferumoxides (Feridex w , tumor contrast 

agent), 5–9 liposomal doxorubicin (Doxil w , chemotherapeutic), 15–17 and albumin-bound paclitaxel 

(Abraxane w , chemotherapeutic), 18–20 to name a few, have all been approved for human clinical 

use in the diagnosis and treatment of a variety of patient tumors and metastatic cancers. For public 

information regarding FDA approval letters, label information, and review packages for each 

of these drugs and others listed in Table 8.1, please visit the following website: http://www. 

accessdata.fda.gov/scripts/cder/drugsatfda/. 

Of the many drugs listed in Table 8.1, tumor contrast agents comprise some of the earliest 

nanosized products to receive FDA approval for human clinical use. Critical to therapeutic 

outcome, diagnostic imaging of nanosized contrast agents using non-invasive magnetic resonance 

imaging (MRI) and positron emission tomography (PET), have shown advanced utility for early 

detection of small tumors, metastatic lymph nodes, and altered tumor microenvironments. 90 

The leading MR contrast agent, Magnevist w , a paramagnetic, gadolinium-based contrast 

medium, and Feridex w , a superparamagnetic iron oxide contrast agent, received primary FDA 

approval for clinical use in 1988 and 1996, respectively. Each contrast agent has a different 

mean particle size (!1 nm and 120–180 nm, respectively), is approved for different clinical indications, 

and demonstrates product-specific adverse sensitivity and pharmacology/toxicology 

profiles in patients. 3–9 

Another early nanomedicine, liposomal encapsulated doxorubicin (Doxil w ), is regulated by the 

FDA and has been available in the clinic for treatment of various cancers since 1995. Drug 

incorporation inside the hydrophilic core or within the hydrophobic phospholipid bilayer coat of 

liposomes, has been shown to improve drug solubility, enhance drug transfer into cells and tissues, 

facilitate organ avoidance, and modify drug release profiles, minimizing toxicity. 91 The liposomal 

formulation of the popular anthracyline, doxorubicin, which is commonly used to treat metastatic 

breast and ovarian cancer, is reported to have diminished cardiotoxicity and enhanced therapeutic 

efficacy compared to the free form of the drug. 15–17 This increased efficacy is most likely due to the 

passive targeting of solid tumors through the enhanced permeability and retention effect inherent to 

tumor vasculature and aberrant tumor morphology. 17 The approximate diameter of the doxorubicin–liposomal 

product is reported to be 100 nm, near the size limits described in the FDA 

definition of a nanomedicine. 

Nanoparticlized, albumin-bound paclitaxel, Abraxane w , is one of the first FDA-approved 

(January, 2005) cancer drugs that was submitted to the FDA with specific claims regarding 

nano-related characteristics that improved this particular formulation over previous submissions. 

18–20 Initially identified as (ABI-007), this chemotherapeutic agent that is approved for 

the treatment of metastatic breast cancer eliminates the excipient Cremophor w by overcoming 

intrinsic insolubility with nanoparticlization of the active paclitaxel ingredient. Although widely 

used in many drug formulations, Cremophor has been associated with an adverse hypersensitivity 

reaction that limited the amount of paclitaxel that could be safely administered. 20 Unlike the 

previous three agents, the approximate diameter of Abraxane is 130 nm, slightly larger than the 


Regulatory Perspective in Nanotechnology Products 143 

requirements described in the National Nanotechnology Initiative (NNI) definition of nanotechnology 

(!100 nm), and thus originally was determined not to fit the size requirement of a 

nanomedicine. However, the fact that the exact particle size exceeds 100 nm may not necessarily 

preclude Abraxane and other similarly-sized or larger products from being 

considered nanotechnology. 

8.3 FDA DEFINITION OF NANOTECHNOLOGY 

As described on the NNI website (http://www.nano.gov/index.html), “Nanotechnology is the 

understanding and control of matter at dimensions of roughly 1–100 nm, where unique phenomena 

enable novel applications. Encompassing nanoscale science, engineering and technology, nanotechnology 

involves imaging, measuring, modeling, and manipulating matter at this length scale.” 

Similarly, in a recent publication by the European Science Foundation on Nanomedicine, 92 nanotechnology 

is defined as the following: “In case of these devices, nanoscale objects were defined as 

molecules or devices within the size range of one to hundreds of nanometers that are the active 

component or object, even within the frame-work of a larger micro-size device or at a macrointerface.” 

A discussion regarding the definition of nanotechnology is currently an active topic 

among various organizations. However, while a definition is always useful, the FDA has traditionally 

dealt with products on a case-by-case basis. This approach is likely to continue for 

nanotechnology products, regardless of the definition. However, we are actively involved in the 

various organizations and committees dealing with issues regarding terminology, such as the 

American Society for Testing and Materials (ASTM) International. 

The definition of nanotechnology at the FDA has been consistent with the definition of nanotechnology 

as reported by the NNI. As such, the FDA currently defines nanotechnology with the 

following definition: 

1. The existence of materials or products at the atomic, molecular, or macromolecular 

levels, where at least one dimension that affects the functional behavior of the drug/device 

product is in the length scale range of approximately 1–100 nm. 

2. The creation and use of structures, devices, and systems that have novel properties and 

functions because of their small size. 

3. The ability to control or manipulate the product on the atomic scale. 

8.4 FDA INITIATIVES IN NANOTECHNOLOGY 

The FDA is responsible for protecting the health of the public by assuring the safety, efficacy and 

security of human and veterinary drugs, biological products, medical devices, our nation’s food 

supply, cosmetics, and products that emit radiation. The FDA is also responsible for advancing the 

public’s health by helping to speed innovations that make medicines and foods more effective, safer 

and more affordable; and by helping the public get the accurate, science-based information they 

need to use medicines and foods to improve their health. Towards this end, the FDA needs to 

proactively pursue those innovations that would result in a significant enhancement of public 

health. Therefore, it falls directly within the mandate of its mission for the FDA to understand 

and help to speed innovations that promise to improve many products that will be regulated by the 

FDA, such as those resulting from nanotechnology. 

In response to a slowdown in innovating medical therapies submitted to the FDA for approval, the 

FDA released the Critical Path Initiative (http://www.fda.gov/oc/initiatives/criticalpath/whitepaper. 

html) in March 2004. The report describes the urgent need to modernize the medical product 

development process—the Critical Path—to make product development more predictable and less 


144 

costly. The Critical Path Initiative at the FDA is a serious attempt to bring attention and focus to the 

need for targeted scientific efforts to modernize the techniques and methods used to evaluate the 

safety, efficacy, and quality of medical products as they move from product selection and design to 

mass manufacture. It is intended to integrate new science into the regulatory process. As such, a list of 

opportunities for the Critical Path Initiative was published by the FDA, and nanotechnology was 

identified as one of the elements under the Critical Path Initiative. Figure 8.1 illustrates the three 

dimensions of the Critical Path (safety, medical utility and industrialization). 

However, the Food and Drug Administration is only one of several government agencies that 

currently regulates and evaluates a wide range of products that utilize nanotechnology and/or 

contain nanomaterials, including foods, cosmetics, drugs, devices, and veterinary products. As a 

member of both the National Science and Technology Council (NSTC) Committee on Technology, 

and the Nanoscale Science, Engineering and Technology (NSET) Working Group on Nanomaterials 

Environmental and Health Implications (NEHI), the FDA coordinates with other government 

agencies, including the National Institute for Environmental Health Sciences (NIEHS), the 

National Toxicology Program (NTP), and the National Institute of Occupational Safety and 

Health (NIOSH) to develop novel strategies for the characterization of and safety evaluation 

standards for nanomaterials. The FDA also works closely with the National Cancer Institute 

(NCI) and the National Institute of Standards and Technology (NIST) in the Interagency Oncology 

Task Force, Nanotechnology subcommittee, to pool knowledge and resources to facilitate the 

development and approval of new cancer drugs and to address the use of nanotechnology in 

cancer therapies and diagnostics. Within the FDA, the Office of Science and Health Coordination 

manages the regular interaction and discussion of nanotechnology between the various agency 

centers and offices, and the nanotechnology working groups within each center, and address the 

concerns about the regulation of nanosized drugs and materials submitted to the individual centers. 

Dimensions 

Basic 

research 

Safety 

Medical 

utility 

Industrialization 

Prototype 

design or 

discovery 

Material selection 

structure 

activity 

relationship 

In vitro and 

computer 

model 

evaluation 

Physical 

design 

Preclinical 

development 

In vitro 

and animal 

testing 

In vitro 

and animal 

models 

Characterization 

small-scale 

production 


Clinical development 

Human 

and animal 

testing 

Human 

efficacy 

evaluation 

Manunfacturing 

scale-up 

refined 

specifications 

FDA filling/ 

approval & 

launch 

preparation 

Safety 

follow 

up 

Mass 

production 

FIGURE 8.1 A highly generalized description of activities that must be successfully completed at different points 

and in different dimensions along the Critical Path. Many of these activities are highly complex—whole industries 

are devoted to supporting them. Not all the described activities are performed for every product, and many activities 

have been omitted for the sake of simplicity. (From http://www.fda.gov/oc/initiatives/criticalpath/whitepaper. 

html.) 



8.5 RESEARCH AT THE FDA ON NANOTECHNOLOGY 

Although the FDA is a regulatory agency, it is also engaged in a number of research activities. 

However, the FDA recognizes that science supported by research provides the foundation for sound 

regulation. As such, the FDA is interested in research projects that can address specific regulatory 

needs. As an example, FDA has a grants program in support of orphan products research and 

development. The FDA does not have a grants program to support other research in non-FDA 

laboratories. Nevertheless, in several of its Centers the FDA conducts research to understand the 

characteristics of nanomaterials and nanotechnology processes. Research interests include any 

areas related to the use of nanoproducts that the FDA needs to consider in the regulation of 

these products. 

Although many FDA Centers are engaged in some nanotechnology research activities, only 

some of the specific projects that are currently ongoing are listed below. The research projects, 

which happen to be conducted by the Center for Drug Evaluation and Research (CDER), include: 

1. Particle size determination in marketed sunscreens with TiO2 and ZnO nanoparticles. 

Sunscreens are considered drugs in the U.S., and although some undergo a premarket 

FDA review, others are marketed under over-the-counter (OTC) monographs that outline 

the active ingredients that the FDA has found to be safe and effective, as well as labeling 

that is truthful and not misleading. TiO2 and ZnO used in sunscreens are therefore 

manufactured according to OTC monographs, and these monographs do not specify 

particle size. As a result, a manufacturer can formulate a sunscreen using particles of 

TiO2 and ZnO that may or may not be in the nano size range. Therefore, to assess the 

particle size of TiO 2 and ZnO in currently marketed sunscreens, CDER has undertaken a 

research project, in collaboration with NIST. 

2. Another project focuses on manufacturing various nanoformulations in our laboratories 

and characterizing the physical and chemical properties of the nanoparticles in these 

formulations, including evaluating the effects of excipients on the measured parameters, 

the effects of preparation methodology on the measured parameters, and the impact (if 

any) that process and formulation variables may have on nanotechnology product 

characteristics and stability. 

3. Another project is studying the in vivo effects of selected nanoparticles, using validated 

animal models. The toxicity of the same nanoparticles will also be evaluated, using 

various in vitro systems. 

8.6 REGULATORY CONSIDERATIONS FOR NANOTECHNOLOGY PRODUCTS 

8.6.1 JURISDICTION OF NANOTECHNOLOGY PRODUCTS 

There has been much interest in the question of whether nanotechnology products will be 

considered drugs, devices, biologics or a combination of the three for regulatory purposes and 

assignment of work. This has been extensively discussed internally within the FDA and the current 

thinking is that many of these products will be considered combination products. 

Combination products (i.e., drug–device, drug–biologic, and device–biologic products) are 

increasingly incorporating cutting edge, novel technologies that hold great promise for advancing 

patient care. For example, innovative drug delivery devices have the potential to make treatments 

safer, more effective, or more convenient to patients. For example, drug-eluting cardiovascular 

stents have the potential to reduce the need for surgery by preventing the restenosis that sometimes 

occurs following stent implantation. Drugs and biologics can be used in combination to potentially 

enhance the safety and/or effectiveness of either product used alone. Biologics are being 


146 


incorporated into novel orthopedic implants to help facilitate the regeneration of bone required to 

permanently stabilize the implants. 

Because combination products involve components that would normally be regulated under 

different types of regulatory authorities, and frequently by different FDA centers, they also raise 

challenging regulatory, policy, and review management issues. A number of criticisms have been 

raised regarding the FDA’s regulation of combination products. These include concerns about the 

consistency, predictability, and transparency of the assignment process; issues related to the 

management of the review process when two (or more) FDA centers have review responsibilities 

for a combination product; lack of clarity about the postmarket regulatory controls applicable to 

combination products; and lack of clarity regarding certain Agency policies, such as when applications 

to more than one agency center are needed. 

To address these concerns, FDA’s Office of Combination Products (OCP) was established on 

December 24, 2002, as required by the Medical Device User Fee and Modernization Act of 2002 

(MDUFMA). The law gives the office broad responsibilities covering the regulatory life cycle of 

drug–device, drug–biologic, and device–biologic combination products. However, the primary 

regulatory responsibilities for, and oversight of, specific combination products will remain in 

one of three product centers—the Center for Drug Evaluation and Research, the Center for 

Biologics Evaluation and Research, or the Center for Devices and Radiological Health—to 

which they are assigned. 

A combination product is assigned to an agency center or alternative organizational component 

that will have primary jurisdiction for its premarket review and regulation. Under section 503(g)(1) 

of the act, assignment to a center with primary jurisdiction, or a lead center, is based on a 

determination of the “primary mode of action” (PMOA) of the combination product. For 

example, if the PMOA of a device–biological combination product is attributable to the biological 

product, the agency component responsible for premarket review of that biological product would 

have primary jurisdiction for the combination product. 

A final rule defining the PMOA of a combination product was published in the August 25, 2005 

edition of the Federal Register, and is available at http://www.fda.gov/oc/combination/. The final 

rule defines PMOA as “the single mode of action of a combination product that provides the most 

important therapeutic action of the combination product.” 

8.6.2 EXISTING FDA REGULATIONS THAT APPLY TO NANOMATERIAL-CONTAINING PRODUCTS 

The FDA has specific guidance/requirements in place for most products. These guidance documents 

can be accessed on the FDA’s website (http://www.fda.gov/oc/industry/default.htm). These 

requirements apply to all products, whether they contain nanomaterials or not. Existing requirements 

are therefore considered to be adequate at this time for most nanotechnology 

pharmaceutical products. 

In the following section, there will be a brief description of the current preclinical requirements 

for the submission of most drugs, to highlight why we believe that our current regulations are 

adequate for the evaluation of the types of nanotechnology products that have been reported as 

being “close” to submission to the agency as an investigational new drug (IND). 

Within CDER, the preclinical requirements for approval to market pharmaceutical products 

include both short-term and long-term toxicity testing in rodent and non-rodent species. Specifically, 

the types of studies conducted by pharmaceutical manufacturers prior to New Drug 

Application (NDA) submission include pharmacology (mechanism of action), safety pharmacology 

(functional studies in various organ systems, most notably the cardiovascular system), absorption, 

distribution, metabolism, and excretion (ADME), genotoxicity (potential to cause mutations in both 

in vivo and in vitro assays), developmental toxicity (to assess effects on reproduction, fertility and 

lactation), irritation studies (to assess local irritation effects), immunotoxicology (to assess effects 

on the components of the immune system), carcinogenicity (to assess the capacity of drugs to 



induce tumors in animal models) and other possible studies (specific studies for a product being 

developed). The current battery of tests, as listed above, is therefore considered to be adequate 

because of the following reasons: the studies use high dose multiples of the drug (usually three 

doses, one that results in low, one that results in medium and one that results in high toxicity), at 

least two animal species are used (usually one rodent and one non-rodent), extensive histopathology 

is conducted on most organs (where most organs are removed, fixed, stained and viewed under the 

microscope to assess if there are structural changes that may have resulted from damage to the 

organs), functional tests are conducted to assess if there are effects on specific organ systems (such 

as the heart, brain, respiratory system, reproductive system, immune system, etc.) and drug treatments 

in animals can be for extended periods of time (up to two years for carcinogenicity studies). 

Additional studies might be requested based on drug-specific considerations and on a caseby-case 

basis. Therefore, as for other drugs, nanotechnology products will be handled on a caseby-case 

basis, depending on the characteristics of the particular product being developed. As 

research provides additional understanding with regards to the toxicity of nanomaterials, the 

FDA may require additional toxicological tests to ensure the safety of the products it regulates. 

8.7 SCIENTIFIC CONSIDERATIONS FOR THE DEVELOPMENT 

OF NANOTECHNOLOGY PRODUCTS 

Ensuring the safety of nanomaterials in drug applications is one of the most important concerns for 

regulators and drug developers. However, as important as safety is, it is also important to 

adequately characterize the nanomaterials used in drug products. Proper characterization and 

manufacturing procedures will ensure that a consistent product is being used, both during preclinical 

and clinical development, and after approval. Some of the questions that have been raised 

internally regarding the assurance of safety of nanomaterial-containing products, as well as questions 

regarding characterization, are listed below. It is not within the scope of this chapter to address 

questions regarding manufacturing and scale-up; however, it should be noted that this is an issue 

that needs serious consideration, because unique characteristics of nanotechnology products are 

likely to require unique manufacturing procedures. 

8.7.1 CHARACTERIZATION OF NANOPARTICLES 

It has become clear to most of those involved in research or development of potential nanotechnology 

applications that it is only possible to assess the safety of nanomaterials when the material 

under study is adequately characterized. If two or more laboratories are working on what they think 

is the same material, then in order for the data from one laboratory to be compared with data from 

another laboratory, there needs to be some confirmation that in fact the nanomaterials being studied 

are actually identical. It has been mentioned by many that for nanomaterials, small differences in 

product characteristics (such as size, surface charge, hydrophobicity, or a number of other attributes) 

may impact product behavior and thus result in vastly different safety profiles. Published data 

show that slight modifications of a nanoparticle have resulted in different toxicity profiles. For 

example, several in vitro studies with Starburst w polyamidoamine (PAMAM) dendrimers and 

Caco-2 cells have shown decreased cytotoxicity of cationic dendrimers upon addition of lauroyl 

or polyethylene glycol (PEG) to the surface of the macromolecules. 93–94 Similar findings of diminished 

toxicity were observed in J774 macrophages upon conjugation of polysaccharides to polymer 

nanoparticles and in fibroblasts and CHO cells with the addition of PEG–silica to the surface of 

CdSe and CdSe/ZnS nanocrystals. 95–96 

Although much less toxicity data is available in vivo, it is evident that surface chemistry is of 

similar importance to the toxico- and pharmacokinetics, toxico- and pharmacodynamics, and 

overall stability of the nanomaterials in vivo. 97–99 When injected into rodents, several generations 

of dendrimers with different surface configurations displayed contrasting organ distribution 


148 


patterns and plasma clearance profiles in vivo, typically depositing within the liver, kidney, and 

pancreas. 98–100 PEG-conjugation to the dendrimer particle surface significantly increased the blood 

half-life and diminished the liver accumulation of dendrimer particles. 99 Similarly, quantum dots 

coated with a complex polymer with carbon alkyl side chains demonstrated greater in vivo stability 

and less genotoxicity than those coated with a simpler polymer or lipid coating. 101 Based on the 

findings of these studies and the collection of pulmonary inhalation data with environmental 

ultrafines and other in vitro cytotoxicity results obtained with different nanomaterials (e.g., 

quantum dots, fullerenes, and nanoshells, etc.), it appears that a thorough characterization of the 

nanomaterial is essential to the interpretation and understanding of toxicity data collected from 

in vitro and in vivo studies. 

To summarize some of the ongoing discussions within CDER’s nanotechnology working 

group, a list of questions has been put together regarding the characterization of nanomaterialcontaining 

products. At this time, it appears that there are no adequate answers to these questions. 

However, it may be that in the future, to satisfy regulatory requirements for nanomaterialcontaining 

products, the questions listed below will need to be addressed by product sponsors. 

1. What are the forms in which particles are presented to the organism, the cells and the 

organelles? Are these particles soluble or insoluble, are the nanomaterials organic or 

inorganic molecules, are the nanomaterials described as nanoemulsions, nanocrystal 

colloid dispersions, liposomes, nanoparticles that are combination products (drug– 

device, drug–biologic, drug–device–biologic), or something else? Does the product 

fall within the currently established definition of nanotechnology? 

2. What are the tools that can be used to characterize the properties of the nanoparticles? 

Can standard tools that are used of other types of products (for example, spectroscopic 

tools) be used to characterize nanomaterial-containing products, or do these products 

require specialized tools that are not widely available for product characterization 

(atomic force microscopy, tunneling electron microscopy, or other specialized modalities)? 

3. Are there validated assays to detect and quantify nanoparticles in the drug product and in 

tissues? 

4. How can long- and short-term stability of nanomaterials be assessed? Specifically, how 

can the stability of nanomaterials be assessed in various environments (buffer, blood, 

plasma, and tissues)? 

5. What are the critical physical and chemical properties of nanomaterials in a drug product 

and how do residual solvents, processing variables, impurities and excipients impact 

these properties? 

6. How do physical characteristics impact product quality and performance? Within what 

range can these physical properties be modified without significantly affecting product 

quality and performance? 

7. What are the critical steps in the scale-up and manufacturing process for nanotechnology 

products? How could manufacturing procedures for nanomaterials differ from those for 

other types of drugs? 

8. How are issues concerning the characterization and manufacturing procedures for nanotechnology 

products assessed, when considering the development of “personalized 

therapies”? In this context, personalized therapies refer to those multifunctional products 

that will be “custom made” for specific patients, depending on issues such as the 

expression of a certain receptor on a tumor or the response of the tumor to a particular 

chemotherapeutic agent. For these types of therapies, what should be the level of 

characterization of the nanomaterials? Does the preclinical testing need to be repeated 

for each modification of the multifunctional molecule, or can there be limited testing to 



cover certain aspects of the product behavior, such as the ADME, the toxicology (a 

bridging study for the acute and/or repeat dose toxicology studies)? What aspects of 

the chemistry, manufacturing and controls (CMC) studies need to be repeated and what 

should be the extent of physical characterization? 

8.7.2 SAFETY CONSIDERATIONS 

Other than characterization, the other major consideration has to do with safety. However, it is only 

within the context of adequately characterized products that safety studies provide value. Therefore, 

although safety is instinctively the primary concern, adequate characterization of 

nanomaterials has to precede any studies on safety. 

As particle size decreases, there may be size-specific effects. These effects may impact safety if 

the materials are more reactive. It has been mentioned in numerous discussions that, as particle size 

decreases, reactivity is expected to rise. The increased reactivity combined with the decreased size 

has led us to pose many questions. Some of these questions are listed below: 

1. Will nanoparticles gain access to tissues and cells that normally would be bypassed by 

larger particles? 

2. If nanoparticles enter cells, what effects do they have on cellular and tissue functions? 

Are the effects transient or permanent? Might there be different effects in different cell 

types, depending on specific tissue targets or receptors? 

3. Once nanoparticles enter tissues, how long do they remain there and where do they 

concentrate? 

4. What are the mechanisms by which nanoparticles are cleared from tissues and blood and 

how can their clearance be measured? 

5. What are some route-specific issues for nanomaterials, and how are these issues different 

than for other types of materials that are not in the nanoscale? 

6. With regards to ADME, specific questions include: 

a. What are the differences in the ADME profile for nanoparticles versus larger particles 

of the same drug? 

b. Are current methods used for measuring drug levels in blood and tissues adequate for 

assessing levels of nanoparticles (appropriateness of method, limits of detection)? 

c. How accurate are mass balance studies, especially if the levels of drug administered 

are very low; i.e., can 100% of the amount of the administered drug be accounted for? 

d. How is clearance of targeted nanoparticles accurately assessed? 

e. If nanoparticles concentrate in a particular tissue, how can clearance be accurately 

determined? 

f. Can nanoparticles be successfully labeled for ADME studies? 


In this chapter, we have tried to highlight some of the steps being taken by the FDA to meet the 

challenges of nanotechnology products. The FDA has engaged in an open discussion with numerous 

groups and organizations, has participated in and sponsored workshops, has established internal 

working groups, has outlined a path for the review of combination nanotechnology products and has 

initiated research projects in nanotechnology. All these efforts have been geared towards a better 

understanding of the science and towards assessing the adequacy of existing regulations for the 

types of products that are expected to be regulated by the FDA. 


150 

At this point, and based on internal discussions, the FDA feels that its current regulations are 

adequate to address the types of nanotechnology products that are being developed as drugs or drug 

delivery systems. The reasons for this have been outlined in this chapter. Although there have been 

requests by investigators and developers of nanotechnology products for the FDA to issue nanotechnology-specific 

guidances, no new guidance documents applicable specifically to 

nanotechnology products will be issued at this time. This is because it is felt that all existing 

Guidance documents apply to these products. All products being developed, whether nanotechnology-based 

or not, are unique and therefore likely to have unique issues requiring consideration. 

As such, and as for all other products reviewed by the FDA, nanotechnology products will be 

reviewed on a case-by-case basis. 

As shown in Figure 8.2, the drug development process is multiphasic and along the path, there 

are numerous opportunities for meetings between the sponsor and the FDA. The first meeting is the 

pre-IND meeting, and it is highly recommended that sponsors of nanotechnology products take the 

opportunity to meet with the FDA at this early stage. An early meeting with the FDA can avoid 

unnecessary or unfocussed studies that could result in a waste of sponsor resources. During early 

meetings, sponsors can present, with minimal data, their preclinical plans to the FDA and obtain 

guidance on how to focus the studies submitted in the IND package. This type of interaction 

between the FDA and the sponsor will be valuable to both parties and will pave the way for the 

most efficient drug development plan. 

Basic 

research 

Industry - FDA 

interactions 

during 

development 

Prototype 

design or 

discovery 

Preclinical 

development 

Pre-IND meeting 

Initial 

IND 

submissions 

Ongoing 

submission 

End of 

phase 

2a meeting 


Clinical development 

Phase 1 Phase 2 Phase 3 

End of phase 

2 meeting 

Pre-BLA or NDA 

meeting 

FDA filing/ 

approval & 

launch 

preparation 

Safety update 

Market 

application 

submission 

IND review phase Application 

review 

phase 

FIGURE 8.2 This figure depicts the extensive industry–FDA interactions that occur during product development, 

using the drug development process as a specific example.* Developers often meet with the agency 

before submitting an investigational new drug (IND) application to discuss early development plans. An IND 

must be filed and cleared by the FDA before human testing can commence in the United States. During the 

clinical phase, there are ongoing submissions of new protocols and results of testing. Developers often request 

additional meetings to get FDA agreement on the methods proposed for evaluation of safety or efficacy, and 

also on manufacturing issues. *Note: Clinical drug development is conventionally divided into three phases. 

This is not the case for medical device development. (From www.fda.gov/oc/initiatives/criticalpath/whitepaper.html.) 



8.9 TRADEMARK LISTING 

1. Magnevist w (Schering Aktiengesellschaft, Berlin, Germany) 

2. Feridex w (Advanced Magnetics, Inc., Cambridge, Massachusetts, U.S.A.) 

3. Rapamune w (Wyeth, Madison, New Jersey, U.S.A.) 

4. Emend w (Merck & CO., Inc., Whitehouse Station, New Jersey, U.S.A.) 

5. Doxil w (Ortho Biotech Products, L.P., Bridgewater, New Jersey, U.S.A.) 

6. Abraxane w (American Pharmaceutical Partners, Inc., Schaumburg, Illinois, U.S.A.) 

7. SilvaGard w (AcryMed, Beaverton, Oregon, U.S.A.) 

8. AmBisome w (Gilead Sciences, Inc., Foster City, California, U.S.A.) 

9. Leunesse w (Nanotherapeutics, Inc., Alachua, Florida, U.S.A.) 

10. VivaGel w (Starpharma Holdings, Ltd., Melbourne, Australia) 

11. Combidex w (Advanced Magnetics, Inc., Cambridge, Massachusetts, U.S.A.) 

12. Cremophor w (BASF Aktiengesellschaft, CO., Ludwigshafen, Germany) 

13. Starburst w (Dendritic Nanotechnologies, Inc., Mount Pleasant, Michigan, U.S.A.) 

REFERENCES 

1. Moghimi, S. M., Hunter, A. C., and Murray, J. C., Nanomedicine: Current status and future 

prospects, FASEB Journal, 19 (311), 330, 2005. 

2. Brigger, I., Dubernet, C., and Couvreur, P., Nanoparticles in cancer therapy and diagnosis, Advanced 

Drug Delivery Review, 54 (5), 631–651, 2002. 

3. Kieffer, S. A., Gadopentetate dimeglumine: Observations on the clinical research process, 

Radiology, 174 (1), 7–8, 1990. 

4. Zhang, H. L., Ersoy, H., and Prince, M. R., Effects of gadopentetate dimeglumine and gadodiamide 

on serum calcium, magnesium, and creatinine measurements, Journal of Magnetic Resonance 

Imaging, 23, 383–387, 2006. 

5. de Lussanet, Q. G., Backes, W. H., Griffioen, A. W., van Engleshoven, J. M., and Beets-Tan, R. G., 

Gadopentenate dimeglumine versus ultra-small superparamagnetic iron oxide for dynamic contrastenhanced 

MR imaging of tumor angiogenesis in human colon carcinoma in mice, Radiology, 229, 

429–438, 2003. 

6. Frank, J. A., Zywicke, H., Jordan, E. K., Mitchell, J., Lewis, B. K., Miller, B., Bryant, L. H. Jr., and 

Bulte, J. W., Magnetic intracellular labeling of mammalian cells by combining (FDA approved) 

superparamagnetic iron oxide MR contrast agents and commonly used transfection agents, Academy 

of Radiology, 9 (Suppl. 2), S484–S487, 2002. 

7. Shapiro, E. M., Skrtic, S., Sharer, K., Hill, J. M., Dunbar, C. E., and Koretsky, A. P., MRI detection of 

single particles for cellular imaging, Proceedings of the National Academy of Sciences of the United 

States of America, 101 (30), 10901–10906, 2004. 

8. Daldrup-Link, H. E., Rudelius, M., Oostendorp, R. A., Settles, M., Piontek, G., Metz, S., 

Rosenbrock, H., Keller, U., Heinzmann, U., Rummeny, E. J., Schlegel, J., and Link, T. M., Targeting 

of hematopoietic progenitor cells with MR contrast agents, Radiology, 228 (3), 760–767, 2003. 

9. Kostura, L., Kraitchman, D. L., Macka, A. M., Pittenger, M. F., and Bulte, J. W., Feridex labeling of 

mesenchymal stem cells inhibits chondrogenesis but not adipogenesis or osteogenesis, NMR in 

Biomedicine, 17 (7), 513–517, 2004. 

10. LaVan, D. A., Lynn, D. M., and Langer, R., Moving smaller in drug discovery and delivery, National 

Review of Drug Discovery, 1, 77–84, 2002. 

11. Miller, J. L., Sirolimus approved with renal transplant indication, American Journal of Health- 

System Pharmacy, 56 (21), 2177–2178, 1999. 

12. Aspeslet, L. J. and Yatscoff, R. W., Requirements for therapeutic drug monitoring of sirolimus, an 

immunosuppressive agents used in renal transplant, Clinical Therapeutics, 22 (Suppl. B), B86–B92, 

2000. 

13. Patel, L. and Lindley, C., Aprepitant-a novel NK1-receptor antagonist, Expert Opinion on Pharmacotherapy, 

4 (12), 2279–2296, 2003. 


152 


14. Wu, Y., Loper, A., Landis, E., Hettrick, L., Novak, L., Lynn, K., Chen, C., Thompson, K., Higgins, 

R., Batra, U., Shelukar, S., Kwei, G., and Storey, D., The role of biopharmaceutics in the development 

of a clinical nanoparticle formulation of MK-0869: A beagle dog model predicts improved 

bioavailability and diminished food effect on absorption in human, International Journal of Pharmaceutics, 

285 (1–2), 135–146, 2004. 

15. Markman, M., Kennedy, A., Webster, K., Peterson, G., Kulp, B., and Belinson, J., Phase 2 trial of 

liposomal doxorubicin (40 mg/m(2)) in platinum/paclitaxel refractory ovarian fallopian tube cancers 

and primary carcinoma of the peritoneum, Gynecologic Oncology, 78, 369–372, 2000. 

16. Frykman, G., Williams, G., and Pazdur, R., Conflicting phase II efficacy date for Doxil, Journal of 

Clinical Oncology, 19 (2), 596–597, 2001. 

17. Laginha, K. M., Verwoert, S., Charrois, G. J. R., and Allen, T. M., Determination of doxorubicin 

levels in whole tumor and tumor nuclei in murine breast cancer tumors, Clinical Cancer Research,11 

(19), 6944–6949, 2005. 

18. Harries, M., Ellis, P., and Harper, P., Nanoparticle albumin-bound paclitaxel for metastatic breast 

cancer, Journal of Clinical Oncology, 23 (31), 7768–7771, 2005. 

19. Nyman, D. W., Campbell, K. J., Hersh, E., Long, K., Richardson, K., Trieu, V., Desai, N., Hawkins, 

M. J., and Von Hoff, D. D., Phase I and pharmacokinetics trial of ABI-007, a novel nanoparticle 

formulation of paclitaxel in patients with advanced nonhematologic malignancies, Journal of 

Clinical Oncology, 23 (31), 7785–7793, 2005. 

20. Sparreboom, A., Scripture, C. D., Trieu, V., Williams, P. J., De, T., Yang, A., Beals, B., Figg, W. D., 

Hawkins, M., and Desai, N., Comparative preclinical and clinical pharmacokinetics of a cremaphorfree 

nanoparticle albumin-bound paclitaxel (ABI-007) and Paclitaxel formulated in cremaphor 

(Taxol), Clinical Cancer Research, 11 (11), 4136–4143, 2005. 

21. Chaw, K. C., Manimaran, M., and Tay, F. E., Role of silver ions in destabilization of intermolecular 

adhesion forces measured by atomic force microscopy in staphylococcus epidermis biofilms, Antimicrobial 

Agents and Chemotherapy, 49 (12), 4853–4859, 2005. 

22. Warriner, R. and Burrell, R., Infection and the chronic wound: A focus on silver, Advanced Skin 

Wound Care, 18 (Suppl. 1), 2–12, 2005. 

23. Baker, C., Pradhan, A., Pakstis, L., Pochan, D. J., and Shah, S. I., Synthesis and antibaceterial 

properties of silver nanoparticles, Journal of Nanoscience and Nanotechnology, 5 (2), 244–249, 

2005. 

24. Takemoto, K., Yamamoto, Y., Ueda, Y., Sumita, Y., Yoshida, K., and Niki, Y., Comparative study 

on the efficacy of Ambisome and Fungizone in a mouse model of pulmonary aspergillosis, Journal of 

Antimicrobial Chemotherapy, 57 (4), 724–731, 2006. 

25. Cetin, H., Yalaz, M., Akisu, M., Hilmioglu, S., Metin, D., and Kultursay, N., The efficacy of two 

different lipid-based amphotericin B in neonatal Candida septicemia, Pediatrics International, 47 

(6), 676–680, 2005. 

26. Barratt, G. and Legrand, P., Comparison of the efficacy and pharmacology of formulations of 

amphotericin B used in treatment of leishmaniasis, Current Opinion in Infectious Diseases, 18 

(6), 527–530, 2005. 

27. The Healthcare Sales and Marketing Network. NanoBio w corporation announces FDA approval for 

phase II clinical trials for herpes labialis treatment. http://salesandmarketingnetwork.com/news_release.php?IDZ14361 

(accessed February 24, 2004). 

28. NewsRX. FDA approves phase II clinical trials for herpes labialis treatment. http://www.newsrx. 

com/article.php?articleIDZ87034 (accessed February 24, 2004). 

29. BioSpace. NanoBio announces FDA approval for phase II clinical trials for herpes labialis treatment. 

http://www.biospace.com/news_story.aspx?StoryIDZ15257520&fullZ1 (accessed February 24, 

2004). 

30. Nanotsunami. NanoBio w completes successful phase 2 herpes trial, prepares for phase 3 trial. http:// 

www.voyle.net/Nano%20Medicine%202005/Medicine%202005-0087.htm (accessed July 22, 

2005). 

31. Abner, S. R., Guenthner, P. C., Guarner, J., Hancock, K. A., Cummins, J. E. Jr., Fink, A., Gilmore, 

G. T., Staley, C., Ward, A., Ali, O., Binderow, S., Cohen, S., Grohskopf, L. A., Paxton, L., Hart, 

C. E., and Dezzutti, C. S., A human colorectal explant culture to evaluate topical microbicides for the 

prevention of HIV infection, Journal of Infectious Diseases, 192 (9), 1545–1556, 2005. 



32. Gong, E., Matthews, B., McCarthy, T., Chu, J., Holan, G., Raff, J., and Sacks, S., Evaluation of 

dendrimer SPL7013, a lead microbicide candidate against herpes simplex virus, Antiviral Research, 

68 (3), 139–146, 2005. 

33. McCarthy, T. D., Karellas, P., Henderson, S. A., Giannis, M., O’Keefe, D. G., Heery, G., Paull, 

J. R. A., Matthews, B. R., and Holan, G., Dendrimers as drugs, discovery and preclinical and clinical 

development of dendrimer-based microbicides for HIV and STI prevention, Molecular Pharmacology, 

2 (4), 312–318, 2005. 

34. Jiang, Y. H., Emau, P., Cairns, J. S., Flanary, L., Morton, W. R., McCarthy, T. D., and Tsai, C. C., 

SPL7013 gel as topical microbicide for prevention of vaginal transmission of SHIV89.6P in macaques, 

AIDS Research and Human Retroviruses, 21 (3), 207–213, 2005. 

35. Ito, I., Ji, L., Tanaka, F., Saito, Y., Gopalan, B., Branch, C. D., Xu, K., Atkinson, E. N., Bekele, B. N., 

Stephens, L. C., Minna, J. D., Roth, J. A., and Ramesh, R., Liposomal vector mediated delivery of the 

3p FUS1 gene demonstrates potent antitumor activity against human lung cancer in vivo, Cancer 

Gene Therapy, 11 (11), 733–739, 2004. 

36. Smalltimes. Study: Nano therapy kills tumors, extends life. http://www.smalltimes.com/document_display.cfm?section_idZ53&document_idZ8429 

(accessed November 11, 2004). 

37. Nanotechnology News. Introgen’s nanoparticle tumor treatment increases survival rate by 70%. 

http://www.azonano.com/news.asp?newsIDZ406 (accessed November 12, 2004). 

38. Nanotsunami. Introgen’s nanoparticle cancer therapy INGN 401 demonstrates promise in the treatment 

of lung cancer. http://www.voyle.net/Nano%20Medicine%202005/Medicine%202005-0052. 

htm (accessed May 17, 2005). 

39. Bourrinet, P., Bengele, H. H., Bonnemain, B., Dencausse, A., Idee, J. M., Jacobs, P. M., and Lewis, 

J. M., Preclinical safety and pharmacokinetic profile of ferumoxtran-10, an ultrasmall superparamagnetic 

iron oxide magnetic resonance contrast agent, Investigative Radiology, 41 (3), 313–324, 

2006. 

40. Saksena, M. A., Saokar, A., and Harisinghani, M. G., Lymphotropic nanoparticle enhanced MR 

imaging (LNMRI) technique for lymph node imaging, European Journal of Radiology, 58 (3), 

367–374, 2006. 

41. Harisinghani, M. G., Saksena, M. A., Hahn, P. F., King, B., Kim, J., Torabi, M. T., and Weissleder, 

R., Ferumoxtran-10-enhanced MR lymphangiography: Does contrast-enhanced imaging alone 

suffice for accurate lymph node characterization?, American Journal of Roentgenology, 186 (1), 

144–148, 2006. 

42. Metz, S., Lohr, S., Settles, M., Beer, A., Woertler, K., Rummeny, E. J., and Daldrup-Link, H. E., 

Ferumoxtran-10-enhanced MR imaging of the bone marrow before and after conditioning therapy in 

patients with non-Hodgkin lymphomas, European Radiology, 16 (3), 598–607, 2006. 

43. Schluep, T., Hwang, J., Cheng, J., Heidel, J. D., Barlett, D. W., Hollister, B., and Davis, M. E., 

Preclinical efficacy of the camptothecin-polymer conjugate IT-101 in multiple cancer models, 

Clinical Cancer Research, 12 (5), 1606–1614, 2006. 

44. Schluep, T., Cheng, J., Khin, K. T., and Davis, M. E., Pharmacokinetics and biodistribution of the 

camptothecin–polymer conjugate IT-101 in rats and tumor-bearing mice, Cancer Chemotherapy and 

Pharmacology, 57, 654–662, 2006. 

45. Newkome, G. R., Yao, Z., Baker, G. R., and Gupta, V. K., Cascade molecules: A new approach to 

micelles. A (27)-Arborol, Journal of Organic Chemistry, 50, 2003–2004, 1985. 

46. Tomalia, D. A., Baker, H., Dewald, J. R., Hall, M., Kallos, G., Maris, S., Roeck, J., Ryder, J., and 

Smith, P. A., New class of polymers: Starburst–dendritic macromolecules, Polymer Journal (Tokyo), 

17, 117–123, 1985. 

47. Wiener, E. C., Brechbiel, M. W., Brothers, H., Magin, R. L., Gansow, O. A., Tomalia, D. A., and 

Lauterbur, P. C., Dendrimer-based metal chelates: A new class of magnetic resonance imaging 

contrast agents, Magnetic Resonance in Medicine, 31 (1), 1–8, 1994. 

48. Grinstaff, M. W., Biodendrimers: New polymeric biomaterials for tissue engineering, Chemistry, 8, 

2839–2846, 2002. 

49. Bourne, N., Stanberry, L. R., Kern, E. R., Holan, G., Matthews, B., and Berstein, D. I., Dendrimers, a 

new class of candidate topic microbicides with activity against herpes simplex virus infestation, 

Antimicrobial Agents in Chemotherapy, 44, 2471–2474, 2000. 


154 


50. Zanini, D. and Roy, R., Practical synthesis of PAMAM athiosialodendrimers for probing multivalent 

carbohydrate–lectin binding properties, Journal of Organic Chemistry, 63, 3486–3491, 1998. 

51. Zhuo, R. X., Du, B., and Lu, Z. R., In vitro release of 5-fluoroacil with cyclic core dendritic polymer, 


52. Liu, M., Kono, K., and Frechet, J. M. J., Water-soluble dendritic unimolecular micelles: Their 

potential as drug delivery agents, Journal of Controlled Release, 65, 121–131, 2000. 

53. Kobayashi, H. and Brechbiel, M. W., Dendrimer-based nanosized MRI contrast agents, Current 

Pharmaeutical Technology, 5, 539–549, 2004. 

54. Konda, S. D., Wang, S., Brechbiel, M., and Wiener, E. C., Biodistribution of a 153 Gd-folate 

dendrimer generationZ4, in mice with folate-receptor positive and negative ovarian tumor xenografts, 

Investigative Radiology, 37, 199–204, 2002. 

55. Hirsch, L. R., Stafford, R. J., Bankson, J. A., Sershen, S. R., Rivers, B., Price, R. E., Hazle, J. D., 

Halas, N. J., and West, J. L., Nanoshell-mediated nar-infared thermal therapy of tumors under 

magnetic resonance guidance, Proceedings of the National Academy of Sciences of the United 

States of America, 100, 13549–13554, 2003. 

56. Loo, C., Lin, A., Hirsch, L., Lee, M. H., Barton, J., Halas, N., West, J., and Drezek, R., Nanoshellenabled 

photonics-based imaging and therapy of cancer, Technology Cancer Research and Treatment, 

3, 33–40, 2004. 


cancer imaging and therapy, Nano Letters, 5, 709–711, 2005. 

58. O’Neal, D. P., Hirsch, L. R., Halas, N., Payne, J. D., and West, J. L., Photo-thermal tumor ablation in 

mice using near-infrared-absorbing nanoparticles, Cancer Letters, 209 (2), 171–176, 2004. 

59. de Jong, W. H., Roszek, B., and Geertsman, R. E., Nanotechnology in medical applications: Possible 

risks for human health. RIVM Report 265001002, Bilthoven, 2005. 

60. Haddon, R. C., Hebard, A. F., Rosseinsky, M. J., Murphy, D. W., Duclos, S. J., Lyons, K. B., Miller, 

B., Rosamilia, J. M., Fleming, R. M., Kortan, A. R., Glarum, S. H., Makjiha, A. V., Muller, A. J., 

Eick, R. H., Zahurak, S. M., Tycko, R., Dabbagh, G., and Thiel, F. A., Conducting films of C60 and 

C70 by alkali-metal doping, Nature, 350, 320–322, 1991. 

61. Bosi, S., Da Ros, T., Spalluto, G., and Prato, M., Fullerene derivatives: An attractive tool for 

biological applications, European Journal of Medical Chemistry, 38, 913–923, 2003. 

62. Friedman, S. H., Decamp, D. L., Sijbesma, R. P., Srdanov, G., Wudl, F., and Kenyon, G. L., 

Inhibition of the HIV-1 protease by fullerene derivatives: Model building studies and experimental 

verification, Journal of the American Chemical Society, 115, 6506–6509, 1993. 

63. Schiniazi, R. F., Sijbesma, R., Srdanov, G., Hill, C. L., and Wudl, F., Synthesis and virucidal activity 

of a water soluble, configurationally stable, derivatized C60 fullerene, Antimicrobial Agents and 

Chemotherapy, 37, 1707–1710, 1993. 

64. Chueh, S. C., Lai, K. M., Lee, M. S., Chiang, L. Y., Ho, T., and Chen, S. C., Decrease of free radical 

level in organ perfusate by a novel water soluble carbon-sixty hexa(sulfobutyl)fullerenes, Transplant 

Proceedings, 31, 1976–1977, 1999. 

65. Lin, A. M., Chyi, B. Y., Wang, S. D., Yu, H., Kanakamma, P. P., Luh, T. Y., Chou, C. K., and Ho, 

L. T., Carboxyfullerene prevents iron-induced oxidative stress in rat brain, Journal of Neurochemistry, 

72, 1634–1640, 1999. 

66. Cagle, D. W., Kennel, S. J., Mirzadeh, S., Alford, J. M., and Wilson, L. J., In vivo studies of 

fullerene-based materials using endohedral metallofullerene radiotracers, Proceedings of the 

National Academy of Sciences of the United States of America, 96, 5182–5187, 1999. 

67. Gharbi, N., Pressac, M., Hadchouel, M., Szwarc, H., Wilson, S. R., and Moussa, F., (60)Fullerene is a 

powerful antioxidant in vivo with no acute or subacute toxicity, Nano Letters, 5 (12), 2578–2585, 

2005. 

68. Bruchez, M. Jr., Moronne, M., Gin, P., Weiss, S., and Alivisatos, A. P., Semiconductor nanocrystals 

as fluorescent biological labels, Science, 281, 2013–2016, 1998. 

69. Sukhanova, A., Venteo, L., Devy, J., Artemyev, M., Oleinikov, V., Pluot, M., and Nabiev, I., Highly 

stable fluorescent nanocrystals as a novel class of labels for immunohistochemical analysis of 

paraffin embedded tissue sections, Laboratory Investigation, 82, 1259–1261, 2002. 



70. Sukhanova, A., Devy, J., Venteo, L., Kaplan, H., Artemyev, M., Oleinikov, V., Klinov, D., Pluot, M., 

Cohen, J. K., and Nabiev, I., Biocompatible fluorescent nanocrystals for immunolabeling of 

membrane proteins and cells, Analytical Biochemistry, 324, 60–67, 2004. 

71. Lidke, D. S., Nagy, P., Heintzmann, R., Arndt-Jovin, D. J., Post, J. N., Freco, H. E., Jares-Erijman, 

E. A., and Jovin, T. M., Quantum dot ligands provide new insights into erbB/HER receptor mediated 

signal transduction, Nature Biotechnology, 22, 198–203, 2004. 

72. Akerman, M. E., Chan, W. C., Laakkonen, P., Bhatia, S. N., and Ruoslahti, E., Nanocrystal targeting 

in vivo, Proceedings of the National Academy of Sciences of the United States of America, 99, 

12617–12621, 2002. 

73. Kim, S., Lim, Y. T., Soltesz, E. G., De Grand, A. M., Lee, J., Nakayama, A., Parker, J. A., 

Mihalijevic, T., Laurence, R. G., Dor, D. M., Cohn, L. H., Bawendi, M. G., and Frangioni, J. V., 

Near-infrared fluorescent type II quantum dots for sentinel lymph node mapping, Nature Biotechnology, 

22, 93–97, 2003. 

74. Ballou, B., Lagerholm, B. C., Ernst, L. A., Bruchez, M. P., and Waggoner, A. S., Noninvasive 

imaging of quantum dots in mice, Bioconjugate Chemistry, 15, 79–86, 2004. 

75. Duncan, R., Nanomedicine gets clinical, Materials Today, 8 (8 Suppl. 1), 16–17, 2005. 

76. Yamawaki, H. and Iwai, N., Cytotoxicity of water-soluble fullerene in vascular endothelial cells, 

American Journal of Physiology. Cell Physiology, 290 (6), C1495–C1502, 2006. 

77. Kamat, J. P., Devasagayam, T. P. A., Priyadarsini, K. I., and Mohan, H., Reactive oxygen species 

mediated membrane damage induced by fullerene derivatives and its possible biological implications, 

Toxicology, 155, 55–61, 2000. 

78. Sayes, C. M., Gobin, A. M., Ausman, K. D., Mendez, J., West, J. L., and Colvin, V. L., Nano-C60 

cytotoxicity is due to lipid peroxidation, Biomaterials, 26, 7587–7595, 2005. 

79. Dugan, L. L., Gabrielsen, J. K., Yu, S. P., Lin, T. S., and Choi, D. W., Buckminsterfullerenol free 

radical scavengers reduce excitotoxic and apoptotic death of cultured cortical neurons, Neurobiology 

of Disease, 3, 129–135, 1996. 

80. Hussain, S. M., Hess, K. L., Gearhart, J. M., Geiss, K. T., and Shlager, J. J., In vitro toxicity of 

nanoparticles in BRL 3A rat liver cells, Toxicology In Vitro, 19, 975–983, 2005. 

81. Jia, F., Wang, H., Yan, L., Wang, X., Pei, R., Yan, T., Zhao, Y., and Guo, X., Cytotoxicity of carbon 

nanomaterials: Single-wall nanotube, multi-wall nanotube, and fullerene, Environmental Science & 

Technology, 39, 1378–1383, 2005. 

82. Renwick, L. C., Donaldson, K., and Clouter, A., Impairment of alveolar macrophage phagocytosis 

by ultrafine particles, Toxicology and Applied Pharmacology, 172, 119–127, 2001. 

83. Moller, W., Hofer, T., Ziesenis, A., Karg, E., and Heyder, J., Ultrafine particles cause cytoskeletal 

dysfunctions in macrophages, Toxicology and Applied Pharmacology, 182, 197–207, 2002. 

84. Oberdorster, G., Toxicology of ultrafine particles: In vivo studies, Philosophical Transactions of the 

Royal Society of London, Series A, 358, 2719–2740, 2000. 

85. Brown, D. M., Wilson, M. R., Macnee, W., Stone, V., and Donaldson, K., Size-dependent proinflammatory 

effects of ultrafine polystyrene particles: A role for surface area and oxidative stress in 

the enhanced activity of ultrafines, Toxicology and Applied Pharmacology, 175, 191–199, 2001. 

86. Johnston, C. H., Finkelstein, J. N., Mercer, P., Corson, N., Gelein, R., and Oberdorster, G., Pulmonary 

effects induced by ultrafine PTFE particles, Toxicology and Applied Pharmacology, 168, 

208–215, 2000. 


evolving from studies of ultrafine particles, Environmental Health Perspectives, 113 (7), 823–839, 

2005. 

88. Holsapple, M. P. and Lehman-McKeeman, L. D., Forum series: Research strategies for safety 

evaluation of nanomaterials, Toxicological Sciences, 87 (2), 315, 2005. 

89. Oberdorster, E., Manufactured nanomaterials (fullerenes, C60) induce oxidative stress in the brain of 

juvenile largemouth bass, Environmental Health Perspectives, 112 (10), 1058–1062, 2004. 

90. Sullivan, D. C. and Ferrari, M., Nanotechnology and tumor imaging: Seizing an opportunity, Molecular 

Imaging, 3 (4), 364–369, 2004. 

91. Portney, N. G. and Ozkan, M., Nano-oncology: Drug delivery, imaging, sensing, Analytical and 

Bioanalytical Chemistry, 384, 620–630, 2006. 


156 


92. Nanomedicine, an ESF-European Medical Research Councils, (EMRC) forward look report. http:// 

www.esf.org/, 2005. 

93. El-Sayed, M., Ginski, M., Rhodes, C., and Ghandehari, H., Transepithelial transport or poly(amidoamine) 

dendrimers across Caco-2 cell monolayers, Journal of Controlled Release, 81, 355–365, 

2002. 

94. Jevprasesphant, R., Penny, J., Jalal, R., Attwood, D., McKeown, N. B., and D’Emanuele, A. D., The 

influence of surface modification on the cytotoxicity of PAMAM dendrimers, International Journal 

of Pharmaceutics, 252, 263–266, 2003. 

95. Lemarchand, C., Gref, R., Pasirani, C., Garcion, E., Petri, B., Muller, R., Costantini, D., and 

Couvreur, P., Influence of polysaccharide coating on the interactions of nanoparticles with biological 

systems, Biomaterials, 27, 108–118, 2006. 

96. Kirchner, C., Liedl, T., Kudera, S., Pellegrino, T., Javier, A. M., Gaub, H. E., Stolzle, S., Fertig, N., 

and Parak, W. J., Cytotoxicity of colloidal CdSe and CdSe/ZnS nanoparticles, Nano Letters, 5 (2), 

331–338, 2005. 

97. Neerman, M. F., Zhang, W., Parrish, A. R., and Simanek, E. E., In vitro and in vivo evaluation of a 

melamine dendrimer as a vehicle for drug delivery, International Journal of Pharmaceutics, 281, 

129–132, 2004. 

98. Nigavekar, S., Sung, L. Y., Llanes, M., El-Jawahri, A., Lawrence, R. S., Becker, C. W., Balogh, L., 

and Khan, M. K., 3H Dendrimer nanoparticle organ/tumor distribution, Pharmaceutical Research, 

21 (3), 476–483, 2004. 


Paulus, W., and Duncan, R., Relationship between structure and biocompatibility in vitro, and 

preliminary studies on the biodistribution of 125I-labeled polyamidoamine dendrimers in vivo, 



(PAMAM) starburst dendrimers, Journal of Biomedical Materials Research, 30 (1), 53–65, 1996. 

101. Hardman, R., A toxicological review of quantum dots: Toxicity depends on physiochemical and 

environmental factors, Environmental Health Perspectives, 114 (2), 165–172, 2006. 


9 

CONTENTS 

Polymeric Conjugates for 

Angiogenesis-Targeted Tumor 

Imaging and Therapy 

Amitava Mitra, Anjan Nan, Bruce R. Line, and 

Hamidreza Ghandehari 

9.1 Introduction ......................................................................................................................... 159 

9.2 Tumor Angiogenesis ........................................................................................................... 161 

9.3 Angiogenesis Markers ........................................................................................................ 162 

9.3.1 VEGF and VEGF Receptors .................................................................................. 162 

9.3.2 Aminopeptidases..................................................................................................... 162 

9.3.3 Matrix Metalloproteinases...................................................................................... 164 

9.3.4 Integrins .................................................................................................................. 166 

9.3.4.1 RGD Multimers ....................................................................................... 167 

9.4 Polymeric Conjugates for Angiogenesis Targeting............................................................ 168 

9.4.1 Polymer-Based Antiangiogenic Gene Therapy ..................................................... 169 

9.4.2 Polymer-Based Nuclear Imaging and Radiotherapy ............................................. 170 

9.5 Summary and Future Directions ......................................................................................... 175 

9.6 List of Abbreviations .......................................................................................................... 175 

Acknowledgment ......................................................................................................................... 177 

References .................................................................................................................................... 177 


Despite significant progress in the surgical treatment, radiotherapy, and chemotherapy of solid 

tumors (e.g., breast, colon, lung, and prostate), too few patients survive long term. This critical 

unmet need has been the long term focus of a wide range of therapies directed at the malignant 

cancer cell phenotype, its dysfunctional apoptotic signaling pathways, and survival adaptations. To 

enhance drug effectiveness and decrease non-specific toxicities, many of these agents have been 

designed to specifically target tumor cell antigens, receptors, and tumor microenvironmental 

markers. 

In contrast to direct targeting of specific tumor cell markers, targeting endothelial cells 

supporting tumor angiogenesis provides an alternative, broadly applicable method for cancer 

diagnosis and therapy. 1–5 All growing tumors require an augmented blood supply, so angiogenesis 

is a critical common denominator for therapeutic attack. Two distinct types of vessel targeted 

therapies have evolved. 6 The first general group, the anti-angiogenic agents, inhibit tumorinduced 

neovascularization by preventing proliferation, migration, and differentiation of 


159

160 

endothelial cells. 7–9 Anti-angiogenic drugs have been carried through substantial development to 

clinical trials. 7,10–12 Bevacuzimab (Avastin), an anti-vascular endothelial growth factor (VEGF) 

monoclonal antibody, has been approved for the treatment of colorectal cancer. 13 The second type, 

the antivascular agents, cause rapid and selective occlusion of existing tumor blood vessels, leading 

to tumor ischemia and extensive hemorrhagic necrosis. 5,14,15 Although still in preclinical development, 

antivascular-targeted therapeutics produce a characteristic pattern of necrosis after 

administration to mice and rats with solid tumors. 16,17 They cause a widespread central necrosis 

that can extend to as much as 95% of the tumor. 16–18 

Indeed, over the past 5–10 years, there has been a rapid development of small molecular 

therapies directed against neovascular angiogenesis. These molecules are typically smaller than 

2 kDa and show very low to absent immunogenicity, high target affinity, rapid targeting, and fast 

blood clearance. Small molecular agents have been directed at a number of angiogenesis-related 

targets and have been coupled to moieties to support scintigraphic, optical, and magnetic resonance 

imaging (MRI). 19,20 Other modifications have supported delivery of drugs, tumor enzymecleaved 

prodrugs, and therapeutic radionuclides. 21 This rich foundation provides the infrastructure 

upon which polymer-based conjugates have allowed further strides. Well-designed 

polymer conjugates demonstrate improved pharmacokinetics with prolonged blood residence, 

lower non-specific tissue penetration, and increased passive tumor localization because of 

enhanced permeability, and retention; i.e., the EPR effect. 22 Polymer conjugates bearing 

targeting ligands often demonstrate high tumor affinity via multivalent interactions. 23 Relative 

to smaller molecular platforms, they can also carry higher diagnostic and therapeutic agent 

payloads (Figure 9.1). 

(a) 

(c) 

(b) 

Biological Payload (Diagnostic or Therapeutic) 


FIGURE 9.1 Schematic of angiogenesis-targeted conjugates. (a) Conjugate of cyclic RGD peptide with a 

therapeutic or a diagnostic agent. Chemotherapeutic or antiangiogenic drugs can be linked via biodegradable 

linkers, whereas imaging agents or radionuclides are generally attached via non-biodegradable linkers. (b) 

Particulate systems such as liposomes and nanoparticles are conjugated to cyclic RGD peptides via a nonbiodegradable 

linker, and the therapeutic or diagnostic agents are encapsulated. (c) Polymeric conjugates have 

cyclic RGD peptides linked onto the side chains via non-biodegradable linker, whereas biodegradable linkers 

can be used for drugs and non-biodegradable linkers for imaging agents or radionuclides. 


Polymeric Conjugates for Angiogenesis-Targeted Tumor Imaging and Therapy 161 

Recent advances provide a clear indication that polyvalent, moderate molecular weight ligandpolymer 

conjugates will provide the platform to target angiogenesis and deliver a broad range of 

effective therapies. This chapter presents an overview of tumor angiogenesis, strategies for 

targeting angiogenic vasculature, and current developments in angiogenesis-targeted polymeric 

conjugates for diagnosis and therapy. 

9.2 TUMOR ANGIOGENESIS 

For tumors to grow and metastasize, they must secure an expanded blood supply by taking over 

existing blood vessels and promoting angiogenesis, i.e., the sprouting of new blood vessels from 

existing vessels. 2,3 Except in wound healing or in the female reproductive tissues, only 0.01% of 

normal endothelial cells are involved in angiogenesis. 24,25 In contrast, angiogenesis is both essential 

for tumor growth beyond 1–2 mm size and is highly specific for neoplasia. 26,27 Indeed, a highly 

significant statistical correlation has been shown between microvessel density and clinical stage, 

histopathology stage, and disease-specific survival. 28,29 

The induction of angiogenesis—the angiogenic switch—is an important early event in tumor 

progression and propagation. 30,31 The switch to an angiogenic phenotype is tightly regulated by a 

balance between endogenous anti-angiogenic 32–34 and pro-angiogenic 35,36 molecules. Studies in 

transgenic mice suggest that the switch occurs early in tumor development as a response to local 

stresses such as hypoxia and low pH. 37,38 Hypoxia increases levels of hypoxia-inducible factors 

(HIF) in tumor cells and in surrounding stromal cells. HIF then promotes the transcription of genes 

for VEGF and several other pro-angiogenic factors. 38,39 

During normal angiogenesis such as during reproduction, development, and wound healing, 

new vessels rapidly mature and stabilize. 2,4 By contrast, tumor blood vessels often fail to 

mature and are characterized by structures that are leaky, tortuous, and irregular in shape. 40 

The high levels of growth factors within a tumor support a continual state of vascular growth, 

regression, and regrowth. Many lack functional pericytes and are exceptionally permeable 

because of the presence of fenestrae, transcellular holes, and a lack of a complete basement 

membrane. 40–44 These ultrastructural changes lead to substantial leaks of tissue fluids into the 

tumor microenvironment and increased interstitial fluid pressure (IFP). Tumor IFP begins to 

increase as soon as the host vessels become leaky in response to angiogenic molecules such as 

VEGF. The importance of this transvascular flux is compounded by the impaired lymphatic 

system characteristic of cancer tissues. Hence, hyperpermeable angiogenic tumor vessels allow 

preferential extravasation of circulating macromolecules, and once in the interstitium, they are 

retained there by a lack of intratumoral lymphatic drainage. This enhanced permeability and 

retention (EPR) effect 22 results in intratumoral penetration and retention of macromolecules. It 

is one important factor explaining the delivery of macromolecular anti-cancer agents to the 

tumor microenvironment. 

In addition to the angiogenic factors that enhance the passive delivery of macromolecules via 

the EPR effect, angiogenic tumor vessels also display a variety of markers (Table 9.1) that can be 

the focus of active molecular targeting. These markers provide several targeting advantages 

relative to direct targeting of tumor cell. 45,46 First, endothelial cells are immediately accessible 

to intravenously delivered ligands. In contrast, the increased interstitial pressure associated with 

the tumor microenvironment can limit the access of macromolecular agents to tumor-associated 

targets. Second, endothelial cells are derived from normal, genetically stable host vessels. Hence, 

they present a target with characteristics that do not fluctuate unlike the genomic instability of 

most cancer tissues. Third, because large numbers of tumor cells rely on each blood vessel, antitumor 

effects will be amplified by vascular bed injury. Endothelial cell damage leads to vascular 

coagulation and downstream hypoxic death of approximately 100-fold more cancer cells. Fourth, 

endothelial targets are common to all solid tumors, whereas most other tumor cell associated 

markers are cancer phenotype specific. 


162 

TABLE 9.1 

Some Ligands and Their Target Receptors with Potential for Targeted Delivery to Tumor 

Angiogenesis 

9.3 ANGIOGENESIS MARKERS 

A relatively large number of endothelial cell surface proteins, ligands, and receptors can distinguish 

tumor angiogenic vessels from normal vasculature (Table 9.1). These markers on tumor endothelial 

cells have been identified by in vitro techniques (e.g., protein electrophoresis, mRNA-based serial 

analysis of gene expression, microarray analysis) 47 and in vivo techniques (e.g., biopanning of 

phage display libraries). 48,49 Many of these targets are the subject of reviews by Ruoshlathi and 

others. 43,44,47,50 

The identification of angiogenesis markers has spurred the development of technologies to 

enhance early cancer diagnosis, therapy planning, and therapeutic intervention. The development 

of diagnostic agents is mainly focused in the areas of nuclear scintigraphy, MRI, and near-infrared 

(NIR) imaging using matrix metalloproteinase (MMP) inhibitors and aVb3 integrin antagonists 

(Table 9.2). 19,20,51,52 Tumor vascular targeting peptides have been used to deliver chemotherapy 

(Table 9.3) given that they selectively target tumor angiogenesis 53,54 and are internalized once 

bound to cell surface receptors. 55,56 The most important classes of angiogenic targets and pertinent 

imaging and therapy studies are briefly discussed below. 

9.3.1 VEGF AND VEGF RECEPTORS 

VEGF has become one of the most important targets for antivascular therapy. It is an endothelial 

cell-specific mitogen that is a potent inducer of angiogenesis. It stimulates endothelial cell growth 

and acts through a family of closely related receptor tyrosine kinases, the most important of which 

is VEGFR-2. 36,57,58 VEGF receptor inhibitors have been shown to have anti-tumor effects. 59,60 The 

most clinically developed is a humanized anti-VEGF antibody (Bevacizumab) that has been 

approved by the U.S. Food and Drug Administration as a first-line therapy for metastatic colorectal 

cancer. 13 

9.3.2 AMINOPEPTIDASES 

Target Targeting Ligand a 

Reference 

aVb3 integrin RGD 53 

Aminopeptidase N (CD13) NGR 54 

Vascular endothelial growth factor Anti-VEGF antibody (Bevacizumab) 13 

aVb3 integrin Anti-aVb3 integrin antibody (Vitaxin) 193 

MMP-2 and MMP-9 HWGF 73 

Aminopeptidase A CPRECESIC 67 

Cell surface nucleolin F3 (N-terminal fragment of human high 

mobility group protein 2) 

194, 195 

Heparan sulfates CGKRK, CDTRL 196 

Unknown SMSIARL 197 

Kallikrein-9 CSRPRRSEC 196 

Platelet-derived growth factor receptor-b CRGRRST 198 

a Single amino acid codes for peptides (see list of abbreviations at end of chapter for details). 


Aminopeptidase N (APN) (also known as CD13) is a membrane, spanning 140 kD cell surface 

protein 61 that plays a role in tumor invasion. 62,63 APN has been successfully targeted using peptides 



TABLE 9.2 

Conjugates for Angiogenesis Targeted Imaging 

Compound Label Target Tumor Model Reference 

Nuclear imaging 

C(RGDfY) I-125 aVb3 integrin Murine osteosarcoma 117 

Gluco-RGD I-125 aVb3 integrin M21 human melanoma 120 

Galacto-RGD F-18 aVb3 integrin M21 human melanoma 121 

c(RGDf(N-Me)V) F-18 aVb3 integrin DLD-1 human colon 

adenocarcinoma 

199 

DOTA-c(RGDyK) Cu-64 aVb3 integrin MDA-MB-435 human 

mammary carcinoma 

118 

HYNIC-RGD4C Tc-99m aVb3 integrin Human renal adenocarcinoma 200 

FB-c(RGDyK)2 F-18 aVb3 integrin U87MG human glioblastoma 136 

DOTA-c(RGDyK)2 Cu-64 aVb3 integrin MDA-MB-435 human 


135 

HYNIC-E-c(RGDfK)2 Tc-99m aVb3 integrin OVCAR-3 human ovarian 

carcinoma 

107 

DOTA-E- c(RGDfK) 2 In-111 aVb3 integrin OVCAR-3 human ovarian 

carcinoma 

107 

DOTA-E{E- c(RGDfK)2}2 Cu-64 aVb3 integrin U87MG human glioblastoma 137 

FB-PEG-c(RGDyK) F-18 aVb3 integrin U87MG human glioblastoma 140 

DOTA-PEG-c(RGDyK) Cu-64 aVb3 integrin U87MG human glioblastoma 141 

DOTA-PEG-E-c(RGDyK)2 Cu-64 aVb3 integrin Human female lung 

adenocarcinoma 

201 

HPMA copolymer-RGD4C Tc-99m aVb3 integrin DU145 & PC-3 human prostate 

carcinoma 

23, 178 

HPMA copolymer-RGDfK In-111 aVb3 integrin Murine lewis lung carcinoma 181 

CGS27023A F-18 MMP Murine ehrlich breast tumor 85, 87 

SAV03M F-18 MMP Murine ehrlich breast tumor 87 

D-Tyr-HWGF I-125 MMP-2, MMP-9 Murine lewis lung carcinoma 83 

DOTA-HWGF 

Magnetic resonance imaging 

Cu-64 MMP-2, MMP-9 MDA-MB-435 human 


90, 91 

Anti-aVb3 antibody (LM609)liposome 

Gadolinium aVb3 integrin Rabbit V2 carcinoma 126 

Anti-aVb3 antibody (DM101)- Gadolinium aVb3 integrin Rabbit corneal micropocket 127 

nanoparticle 

model 

aVb3 peptidomimetic antagonistnanoparticle 

Optical imaging 

Gadolinium aVb3 integrin Rabbit V2 carcinoma 110 

Cy5.5-c(RGDyK) Cy5.5(NIR dye) aVb3 integrin U87MG human glioblastoma 111 

Cy5.5- E- c(RGDyK) 2 Cy5.5 aVb3 integrin U87MG human glioblastoma 129 

Cy5.5-E{E- c(RGDfK)2}2 Cy5.5 aVb3 integrin U87MG human glioblastoma 129 

Cy5.5-GPLGVRGK-PEG-PLL Cy5.5 MMP-2 HT1080 human fibrosarcoma 94 

NIRQ820-GVPLSLTMGC-Cy5.5 Cy5.5 MMP-7 HT1080 human fibrosarcoma 97 

containing an Asn-Gly-Arg (NGR) motif. 54 NGR peptides have been used to deliver chemotherapeutic 

agents 64 and pro-apoptotic peptides 65 to tumor cells. In a murine breast carcinoma xenograft 

model, a NGR-doxorubicin conjugate significantly decreased metastasis to lymph nodes and 

substantially increased survival compared to free doxorubicin. In addition, the conjugates were 

less toxic to the liver and heart as compared to the free drug. 64 


164 

Aminopeptidase A (APA) is a homodimeric membrane, spanning cell surface protein that is 

upregulated in the pericytes of tumor blood vessels. 66 Phage display studies have identified an APA 

binding peptide (CPRECESIC) that inhibits enzymatic activity of APA, suppresses both migration 

and proliferation of endothelial cells, and inhibits tumor growth. 67 

9.3.3 MATRIX METALLOPROTEINASES 

MMPs are a family of over 21 zinc-dependent neutral endopeptidases that play an important role in 

tumor angiogenesis, tissue remodeling, and cell migration. 68–70 The two most important MMPs in 

cancer progression are MMP-2 (gelatinase A) and MMP-9 (gelatinase B). 71,72 In cancer, levels of 

these MMPs are abnormally elevated, enabling cancer cells to degrade the extracellular matrix 

(ECM), invade the vascular basement membrane, and metastasize to distant sites. 70 

MMP targeting for tumor imaging or therapy can be achieved using small molecule inhibitors 

of MMPs, peptide sequences that target MMP, tissue inhibitors of MMPs, and MMP peptide 

substrates that are cleaved at the tumor site. For example, cyclic peptides containing a HWGF 

motif can specifically inhibit MMP-2 and MMP-9 activities and prevent both tumor growth and 

metastasis. 73 Several small molecule MMP inhibitors such as the sulfonamide–hydroxamate 

compound CGS 27023A are currently in clinical trials. 69,74–76 

The protease functionality of MMP provides additional opportunities useful in enzyme prodrug 

therapy (Table 9.3). For example, drugs such as doxorubicin, auristatins, and duocarmycin may be 

conjugated to an acetyl-PLGL peptide sequence, MMP substrate, such that when incubated with 

MMP-2 or MMP-9, cleavage at the GL bond liberates a leucyl drug. Kline et al. showed that such a 

strategy produced no other cleavage intermediates and that proteolysis products were equipotent 

with the parent drugs. 77 

MMP-substrate-doxorubicin conjugates are found to be preferentially metabolized in murine 

fibrosarcoma xenografts relative to plasma and cardiac tissues. 78 There was a 10-fold increase in 

TABLE 9.3 

Conjugates for Angiogenesis Targeted Drug Delivery 


Compound Drug Target Tumor Model Reference 

NGR-doxorubicin Doxorubicin Aminopeptidase N MDA-MB-435 human mammary 

carcinoma 

64 

mPEG-GPLGVdoxorubicin 

Doxorubicin MMP-2, MMP-9 Murine lewis lung carcinoma 79 

Albumin- GPLGIAGQdoxorubicin 

Doxorubicin MMP-2, MMP-9 A375 human melanoma 82 

PLGL-doxorubicin Doxorubicin MMP-2, MMP-9 HT1080 human fibrosarcoma 78 

RGD4C-doxorubicin Doxorubicin aVb3 integrin MDA-MB-435 human mammary 

carcinoma 

64 

RGD4C-doxorubicin Doxorubicin aVb3 integrin MH134 murine hepatoma 132 

RGD4C-AFK-doxorubicin Doxorubicin aVb3 integrin HT1080 human fibrosarcoma 130, 131 

E-c(RGDyK) 2-paclitaxel Paclitaxel aVb3 integrin MDA-MB-435 human mammary 

carcinoma 

202 

PTK787-albumin-PEG- PTK787 (kinase aVb3 integrin Human umbilical vein endothelial cells 21 

c(RGDfK) 

inhibitor) 

c(RGDfC)-PEG-liposomes 5-FU aVb3 integrin Murine B16F10 melanoma 185 

RGD-PEG-liposomes Doxorubicin aVb3 integrin Murine B16F10 melanoma 203 

c(RGDfC)-PEG- Doxorubicin aVb3 integrin Cl-66 murine metastatic mammary 186 

nanoparticles 

tumor carcinoma 



doxorubicin tumor-to-heart ratio and a greater effectiveness than free doxorubicin at reducing 

tumor growth. In particular, the conjugate cured 8 of 10 mice at levels below the maximum 

tolerated dose, whereas doxorubicin cured 2 of 20 mice at its maximum tolerated dose. Furthermore, 

mice treated with the conjugate had no detectable change in body weight or circulating 

reticulocytes. 78 

Other studies have focused on molecular modifications to modify in vivo biodistribution 

kinetics. For example, MMP substrate peptide–doxorubicin conjugates have been coupled to polyethylene 

glycol (PEG) to increase blood circulation time and to reduce non-specific cytotoxicity. 79 

The anti-cancer drug conjugate, mPEG–GPLGV–DOX, was synthesized by conjugating doxorubicin 

with GPLGV peptide (a substrate for MMP-2 and MMP-9) and PEG methyl ether (mPEG). 

In vivo experiments showed that a 50 mg/kg dose of mPEG–GPLGV–DOX had similar therapeutic 

effectiveness as a 10 mg/kg dose of doxorubicin, but it was associated with lower toxicity and 

prolonged life span relative to free drug. 79 

An interesting strategy to deliver drug therapy by EPR was investigated using a maleimide– 

doxorubicin conjugate of an octapeptide MMP substrate (GPLGIAGQ). 80 The water-soluble 

substrate was designed to form an in situ complex with an albumin thiol (cysteine-34). 81 The 

conjugate was efficiently cleaved by activated MMP-2 and MMP-9, liberating a doxorubicin 

tetrapeptide that showed antiproliferative activity in vitro in a murine renal cell carcinoma. 

When tested in a murine model of a human melanoma xenograft, the doxorubicin–octapeptide 

conjugate showed complete binding with albumin within 5 min and a 16 h stability half-life. 82 The 

maximum tolerated dose of the conjugate was 4-fold greater than free doxorubicin. Furthermore, 

the conjugate showed superior anti-tumor effects as compared to free doxorubicin with duration of 

remission for up to 40 days and almost a 4-fold reduction in tumor size as compared to free 

doxorubicin at the end of the 50 day trial period. 

Although imaging MMP expression is a relatively new area of research, a number of radiolabeled 

imaging agents have been tested in animal models (Table 9.2). 19,20 I-125 labeled cyclic 

decapeptides containing HWGF have shown relatively disappointing results in mice bearing Lewis 

lung carcinoma. 83 There was low to moderate uptake in tumor, significant tumor washout overtime, 

high concentrations in the liver and kidney, and low metabolic stability of the iodo-tyrosine 

complex. 83 More promising results have been obtained with 18 F and 11 C labeled MMP inhibitors 

developed for positron emission tomography (PET) imaging. 84–86 In vitro studies showed no loss of 

MMP inhibitory activity of the radiolabeled compound as compared to the parent compound. 84,87 

Dynamic microPET images of tumor-bearing mice showed substantially higher tumor uptake than 

other background organs. 88,89 The cyclic HWGF peptide has been conjugated with a 1,4,7,10-tetraazacylcododecane-N,N 

0 ,N 00 ,N 000 -tetraacetic acid (DOTA) chelate to enable 64 Cu PET imaging. 90,91 

Studies of the 64 Cu labeled peptide in a murine breast cancer model showed improved pharmacokinetics, 

tumor accumulation, and metabolic stability relative to a comparable 125 I 

radiolabeled peptide. 

111 In-diethylenetriaminepentaacetic acid (DTPA)-N-TIMP-2, a radiolabeled high affinity 

endogenous tissue inhibitor of MMP, has been found to have similar activity as unmodified 

N-TIMP-2. 92 Unfortunately, 111 In-DTPA-N-TIMP-2 yielded disappointing results with insignificant 

tumor uptake in clinical studies in patients with Kaposi’s sarcoma. 93 

Optical imaging methods to assess MMP activity has been achieved using a NIR fluorescence 

probe attached to a MMP-2 substrate. 94–96 The MMP-2 imaging probe consists of three elements: a 

quenched NIR fluorochrome (Cy5.5), a MMP-2 peptide substrate (GPLGVRGK), and a PEG-poly- 

L-lysine (PEG–PLL) graft copolymer. 94 The conjugate is designed such that the peptides are 

cleaved by MMP-2 at the tumor site, resulting in several hundred percent increases in the NIR 

fluorescence, allowing visualization of the MMP-2 in human fibrosarcoma tumor. Next MMP-2 

inhibition was imaged after treatment with prinomastat where the treated tumor showed significantly 

lower NIR signal as compared to the untreated tumors. 94 A similar NIR fluorescencebased 

probe has been designed for MMP-7 that is over expressed in colon, breast, and pancreatic 


166 

cancer. 97 In tumor extract containing MMP-7 in vitro, the fluorescence signal from the probe was 

enhanced 7-fold after enzymatic degradation. This probe could also be used to image tumorassociated 

enzyme activity and also inhibition of MMP-7 in vivo. 

9.3.4 INTEGRINS 


Integrins are a diverse group of heterodimeric cell surface receptors for ECM molecules. 98,99 This 

group consists of at least 25 integrins through the pairing 18 a and 8 b subunits. They mediate cell 

adhesion, 100 help in cell migration during metastasis, 101 and regulate cell growth. 102 a Vb 3, a Vb 5, 

and a 5b 1 integrins are upregulated in angiogenic endothelial cells and are essential for angiogenesis. 

103–106 

High affinity aVb3 (and aVb5) selective ligands containing the tripeptide Arg-Gly-Asp (RGD) 

have been identified by phage display studies 53 and have been used to deliver chemotherapeutic 

agents, 64 pro-apoptotic peptides, 65 and radiotherapy 107,108 to tumor tissues. RGD peptides and 

RGD mimetics 109 have been used to detect cancer when coupled to image contrast agents designed 

for MRI, 110 gamma scintigraphy, 19 PET, 19,51 and NIR fluorescence imaging. 111 

A large number of radiolabeled tracers based on the RGD tripeptide sequence have been 

developed (Table 9.2). 19,51,52 Structure activity relationship investigations of the cyclic pentapeptide, 

cyclo-(Arg 1 -Gly 2 -Asp 3 -D-Phe 4 -Val 5 ), 112–114 revealed that a hydrophobic amino acid in 

position 4 is essential for high aVb3 affinity and that the amino acid in position 5 could be modified 

without loss of activity. 115 A tyrosine residue at position 4 or 5 provides a site to radio-iodinate the 

peptide, and lysine residue introduced at position 5 enables 18 F labeling. 116,117 Besides radiohalogenation, 

cyclic RGD peptides can be labeled with radio-metals by introducing a lysine residue at 

the terminal end of the peptide and conjugating that with a chelator such as DOTA, 118 DTPA, 119 

hydrazinonicotinamide (HYNIC), 107 or N-u-bis(2-pyridylmethyl)-L-lysine (DPK). 23 Other radiometals 

such as yttrium-90 ( 90 Y) may provide a means to deliver molecularly targeted radiotherapy. 

For example, in a murine ovarian carcinoma xenograft model, a single injection of a 90 Y labeled, 

dimeric-RGD peptide-DOTA conjugate at its maximum tolerated dose (37 MBq) caused significant 

tumor growth delay and increased median survival time. 107 

Unfortunately, many of these tracers are lipophilic and are both concentrated and excreted by 

the liver. 117 Non-tumor bearing organs that retain the tracers may receive an unacceptably high 

radiation dose. To improve their pharmacokinetics, glucose or a galactose-based sugar amino acid 

has been introduced into the pentapeptide architecture. The resulting iodo-gluco-RGD 120 or [ 18 F]galacto-RGD 

121,122 shows reduced concentration in the liver and increased accumulation in 

tumor. In other attempts to improve pharmacokinetics, cyclic-RGD peptides have been conjugated 

to tetrapeptides of D-amino acids. 123 The pharmacokinetics and tumor accumulation of the 

resulting 18 F labeled conjugates containing D-asparatic acid are found to be comparable to [ 18 F]galacto-RGD. 

123 

Non-invasive imaging of angiogenesis markers should be important, both in therapy planning 

and in monitoring anti-angiogenic therapy. The first human images of aVb3 expression have been 

obtained with [ 18 F]-galacto-RGD in patients with malignant melanoma and sarcoma. 124 These 

biodistribution studies demonstrated highly favorable pharmacokinetic profiles with specific 

receptor binding, rapid blood clearance primarily through the kidney, increasing tumor-to-background 

(T/B) ratios with time, and almost 4-fold higher tumor distribution volume as compared to 

muscle. 125 Tumor aVb3 expression was imaged with good contrast in most parts of the body 

with the exception of the region near the urogenital tract, the tracer excretion route. 

Detection of tumor angiogenesis by aVb3 targeted MRI has been achieved using liposomes or 

nanoparticles encapsulating gadolinium ion that was derivatized with anti-aVb3 antibody 

LM609, 126 antibody DM101 127 or RGD mimetics. 110,128 This approach provided detailed images 

of the tumor and delineation of the tumor from surrounding tissues, and it highlighted angiogenic 



vasculature. This enhanced detection of tumors may be helpful in directing therapeutic 

interventions. 

Several optical angiogenesis imaging methods using cyclic RGD peptides have been described. 

One such method utilized a NIR dye Cy5.5 conjugated to RGD to visualize glioblastoma xenograft 

in nude mice. 111 The tumor could be visualized with high contrast to the background organs as early 

as 30 min post-injection. This study showed the feasibility of in vivo optical imaging of aVb3 

expressions. In a subsequent study, RGD mono-, di- and tetramer conjugated to Cy5.5 were 

compared in a glioblastoma xenograft model. 129 The tumor xenograft could be visualized with 

all three probes 30 min post-injection. The tumor uptake was in the order tetramerOdimerO 

monomer; however, the increase in the T/B ratios for the multimers was modest. For example, 

the dimer and the tetramer showed lower tumor-to-kidney ratios than the monomer because of 

higher net positive charge. Therefore, only modest enhancement of therapeutic index is possible by 

multimerization of RGD peptides and further optimization of these probes is needed. 

RGD-targeted prodrugs have been tested as a means to improve chemotherapy toxicity profiles 

(Table 9.3). One strategy utilizes an enzymatically cleavable tripeptide sequence (D-Ala-Phe-Lys) 

recognized by the tumor-associated protease, plasmin, as a linker between RGD4C and doxorubicin. 

130 Upon incubation with plasmin in vitro, doxorubicin is released and regains its cytotoxicity 

for endothelial and fibrosarcoma cells. Toxicity studies in BALB/c and nude mice showed significantly 

higher weight loss and mortality for free doxorubicin relative to an equimolar dose of the 

prodrug. 131 Furthermore, in vivo evaluation in mouse breast cancer and human adenocarcinoma 

models showed that prodrug anti-tumor efficacy was associated with strong inhibition of angiogenesis. 

131 A similar RGD4C-doxorubicin conjugate was tested in an a Vb 3 negative murine hepatoma 

model and was found to have superior anti-tumor effectiveness as compared to free doxorubicin. 132 

9.3.4.1 RGD Multimers 

As experience with RGD-targeted compounds has evolved, efforts to enhance effectiveness have 

focused on improving tumor uptake and reducing non-specific localization in normal tissues. To 

enhance tumor targeting, RGD multimers have been studied as means to increase receptor binding 

affinity through polyvalency. A dimeric RGD peptide labeled with 111 In, 107 99m Tc, 133,134 64 Cu, 135 

and 18 F 136 have all shown higher receptor binding affinity and significantly higher tumor accumulation 

than the RGD monomer. Similarly, tetrameric RGD peptides have shown almost 3-fold and 

10-fold higher integrin affinity than the dimer and monomer, respectively. 129,137 Biodistribution 

studies in M21 melanoma xenograft-bearing mice indicated that tumor uptake and T/B ratios 

increased as monomer!dimer!tetramer. 138 However, more recent biodistribution studies in 

human glioblastoma-bearing nude mice revealed that the higher tumor uptake of the tetramer as 

compared to dimer was offset by increased liver and kidney uptake. 129,137 This significantly 

lowered the T/B ratios of the multimers, substantially reducing the therapeutic index. 

To address the unfavorable pharmacokinetics of these cyclic peptides, investigators have 

studied their conjugation to hydrophilic polymers such as PEG. A comparative biodistribution 

study showed that free peptide had a more rapid tumor washout than a pegylated cyclic RGD 

conjugate. In contrast, the PEG–RGD conjugate showed a gradual increase in both tumor accumulation 

and tumor-to-blood ratios. 139 Similarly, studies of 18 F 140 and 64 Cu 141 radiolabeled pegylated 

RGD have yielded high quality microPET images of glioblastoma xenografts with higher T/B ratios 

than the non-pegylated peptide. 

As might be expected, studies combining multivalency and pegylation have shown significantly 

improved tumor retention and tumor-to-normal tissue distribution ratios. 142 The tumor uptake of 

pegylated RGD multimers increased by about 3-fold from monomer to tetramer. Tetramer tumorto-liver 

ratios were 20- and 3-fold higher than the monomer and dimer, respectively. 142 


168 

9.4 POLYMERIC CONJUGATES FOR ANGIOGENESIS TARGETING 

Polymeric conjugates have been used for targeted delivery of drugs to tumor sites 143–146 because 

the attachment of drugs to water soluble polymers increases their aqueous solubility, reduces side 

effect, and overcomes multi-drug resistance; the large size of the conjugates increases the blood 

half-life and significantly alters the drug properties and pharmacokinetics; the conjugates can be 

tailor-made (i.e., side-chain content, molecular weight, charge, etc.) for specific targeting and 

delivery needs; they can be designed to passively (EPR) or actively target tumor sites; and sitespecific 

drug release can be achieved by designing biodegradable spacers that can be enzymatically 

cleaved or that are pH sensitive. These advantages have led to the development of a wide range of 

polymer-anti-cancer drug conjugates, some of which are currently in clinical trials. 145 

N-(2-hydroxypropyl) methacrylamide (HPMA) copolymers have shown promise as drug 

carriers. 144,145,147,148 HPMA copolymers have been employed to modify the in vivo biodistribution 

of chemotherapeutic agents and enzymes. The advantages of HPMA copolymers over other watersoluble 

polymers is that they can be tailor-made with simple chemical modifications to regulate 

drug and targeting moiety content for biorecognition, internalization, or subcellular trafficking 

depending on specific therapeutic needs (Figure 9.2). 144,149,150 The overall molecular weight of 

HPMA copolymers is determined by the polymerization conditions, particularly the concentration 

of initiator and chain transfer agents. 151 Various side chain moieties (isotope chelators, targeting 

moieties, and drugs) (Figure 9.2) may be directly linked to the polymer chain via a biodegradable or 

non-biodegradable spacer. 143,144 

Polymer-based delivery systems have been used as carriers for passive and active targeting of 

drugs in the treatment of various diseases and as novel imaging agents. 145,152 Without a specific 

targeting ligand, moderate-sized (O30 kD) polymers can passively (via EPR) accumulate in tumor 

tissues. 22,153 The EPR effect has been used to deliver macromolecular bioactive agents to solid 

tumors, including anti-angiogenic drugs. 154,155 

There are a number of important differences between small molecular weight drugs and 

polymeric conjugates of these small molecules. The advantage of polymer-based delivery 

1 2 3 4 5 

H CH 3 

2 

( C C ) 

x 

C O 

CH 3 

H 2 

( C C ) 

y 

C O 

H CH 3 

2 

( C C ) 

m 

C O 

H 

CH 3 

2 

(C C ) 

n 

C O 

NH 

NH O 

NH 

NH 

CH 2 HC C NH 2 CH 2 

CH 2 

HC OH CH 2 

C O 

C O 

CH 3 

NH 

NH 

OH 

CH 2 

C O 

NH HOOC 

CH 2 

C O 

H 

C NH 

Peptide CH 2 

RGD4C, RGDfK 

CH 2 

CH 2 

CH 2 

H N 

2C 2C CH 2 

125 I, 123 I, 124 I, 131 I 

N 

188 Re, 99m Tc 

N 

H CH 

2 

3 

( C C ) 

z 

C O 

NH 

CH 2 

CH 2 

CH 2 

NH 

C S 

NH 

COOH 

N 

COOH 

N 

COOH 

N 

COOH 

COOH 


177 Lu, 90 Y, 213 Bi, 210 Po 

FIGURE 9.2 Structure of water-soluble HPMA copolymers for angiogenesis targeted delivery of radionuclides 

for imaging and therapy. The conjugate can contain side-chains of different chelating comonomers 

for a wide range of radionuclides for use as a single agent for both diagnosis and therapy. 1-HPMA, 2-MA-Tyr, 

3-MA-GG-RGD4C, 4-MA-GG-DPK, and 5-APMA-CHX-A 00 -DTPA (see list of abbreviations for details). 

(From Mitra, A., et al., Nucl. Med Biol., 33, 43, 2006. With permission.) 



systems stem from decreased extravasation in normal tissues because of the large molecular weight 

of the conjugates. 156 Low extravasation of polymer conjugates in normal tissues generally results in 

reduced systemic toxicity. In this regard, the predominant liver 117 and kidney 120 uptake of small 

RGD peptides has been identified as a significant disadvantage of targeting tumor angiogenesis with 

small peptides. 

The polymeric backbone may also provide a platform to support multivalent targeting ligands 

(Figure 9.1). As previously discussed, conjugation of a targeting peptide onto a polymer backbone 

significantly enhances the tumor to normal tissue uptake in comparison to the peptide itself. This is 

likely attributable to an increased target affinity because of the multivalency of the targeting moiety 

on the polymer backbone, 157 a combination of active targeting and passive EPR effect of the 

macromolecular conjugate, 158 and a decreased extravasation in normal tissues as a result of the 

large molecular weight of the conjugates. 156 

The strategy of targeting the tumor angiogenic endothelium using polymeric conjugates is 

particularly attractive where angiogenesis inhibitors show substantial toxicity at the effective 

dose, e.g., the fumagillin analogue TNP-470 shows extreme neurotoxicity at this level. 159 Conjugation 

of TNP-470 to a water-soluble polymer via an enzymatically degradable spacer allows 

passive localization and drug release at the tumor site and prevents blood–brain barrier 

penetration. 154 

9.4.1 POLYMER-BASED ANTI-ANGIOGENIC GENE THERAPY 

Anti-angiogenic gene therapy strategies against several human tumors have shown encouraging 

results in animal models. 160,161 A number of strategies to deliver therapeutic genes to angiogenic 

endothelial cells using RGD guided viral 162,163 and non-viral systems (cationic polymers, 164 

cationic lipids, 165 and cationic peptides 166 ) have been investigated. The cationic polymer polyethylenimine 

(PEI) has been extensively used as a non-viral gene carrier. 167 PEI has also been 

investigated as an RGD guided transfection system for angiogenic endothelial cells (Table 9.4). 

Broadly, these systems can be divided into two classes: those that are direct conjugates of 

PEI–RGD 168,169 and those where RGD is conjugated onto PEI via a hydrophilic PEG spacer 

(PEI–PEG–RGD). 164,168,170–172 Because of their reduced non-specific interaction with normal 

tissues, the charge-shielded pegylated systems have shown more efficient targeting and gene 

transfection. 171,172 

The PEI–PEG–RGD conjugate architecture has been used to deliver siRNA as a means to 

inhibit VEGFR-2 expression. 170 Specific tumor accumulation was observed in a murine neuroblastoma 

model, and there was lower non-specific uptake in lung and liver as compared to aqueous 

siRNA and non-targeted conjugate. With a 40 mg treatment dose repeated every three days, there 

TABLE 9.4 

Conjugates for Angiogenesis Targeted Gene Therapy 

Compound Gene Target Tumor Model Reference 

c(RGD)-PEG-PEI siRNA aVb3 integrin N2A neuroblastoma 170 

RGD4C-PEG-PEI sFlt-1 aVb3 integrin Human skin capillary endothelial cells 164 

Pronectin F Cw 

LacZ aVb3 integrin Murine meth-AR-1 fibrosarcoma 173 

RGDC-PEG-PEI Luciferase aVb3 integrin Mewo human melanoma 168 

RGD-PEI Plasmid DNA aVb3 integrin Human cervical carcinoma 169 

RGD-PEG-PEI DNA aVb3 integrin Not reported 172 

RGD-lipid nanoparticles Raf-1 aVb3 integrin CT26 human colon carcinoma 165 

RGD4C-PEG-cholesterol liposome CAT aVb3 integrin PO2 human colon carcinoma 204 


170 

was significant inhibition of tumor angiogenesis, reduced tumor growth rate, and a marked 

reduction of peritumoral vascularization. 170 In another study, Kim et al. reported an anti-angiogenesis 

strategy to block VEGF receptor function and thereby inhibit endothelial cell proliferation. 164 

The authors conjugated PEI–PEG–RGD and a sFlt-1 gene that encoded a soluble fragment of Flt-1 

(VEGF receptor antagonist) and found that this non-viral gene delivery system enabled stable 

expression of sFlt-1 by endothelial cells. The expressed soluble Flt-1 fragment, coupled to exogenous 

VEGF, blocked Flt-1 receptor binding and inhibited cultured endothelial cell proliferation. 164 

Using another approach, Hosseinkhani and Tabata investigated the feasibility of tumor targeting 

using a non-viral gene carrier synthesized from a genetically engineered silk-like polymer 

containing repeated RGD sequences(Pronectin FC). 173 When cationized, Pronectin FC-plasmid 

DNA complexes with or without pegylation were intravenously injected into mice carrying a 

subcutaneous Meth-AR-1 fibrosarcoma mass, and the level of gene expression in the tumor was 

significantly higher for the PEG–DNA complex relative to non-pegylated complexes and free 

plasmid DNA. 

Although anti-angiogenic gene therapy is still in its infancy, the studies beginning to appear in 

the literature suggest the prospects of development of new polymer-based gene delivery systems 

with improved transfection efficiency and reduced toxicity. It appears that many of the targeting and 

biodistribution lessons learned in studies of polymer-based drug delivery to sites of angiogenesis 

may be applied to enhance the success of polymer based gene delivery as well. 

9.4.2 POLYMER-BASED NUCLEAR IMAGING AND RADIOTHERAPY 

With the appropriate delivery system, radioisotopes have a significant advantage over other therapy 

agents, namely, the emission of energy that can kill at a distance from the point of radioisotope 

localization. This diameter of effectiveness helps to overcome the problem of tumor heterogeneity 

because, unlike other molecular therapy (cell toxins, chemotherapy, etc.), not all tumor cells need to 

take up the radioisotope to eradicate a tumor. There are also physical characteristics (type of 

particle emission, emission energy, half-life) of different radioisotopes that may be selected to 

enhance therapeutic effectiveness. 174 For example, different isotopes deliver beta particulate ionization 

over millimeters ( 131 I) to centimeters ( 90 Y). Long-lived isotopes such as 131 I that remain 

within the tumor target may provide extended radiation exposure and high radiation dose, 

especially if there is progressive renal clearance and high target to non-target ratios. 

RGD peptides labeled with therapeutically relevant isotopes such as b-particle emitters have 

been investigated as potential angiogenesis targeted radiotherapy (Table 9.5). The chelation conditions 

for 90 Y and lutetium-177 ( 177 Lu) labeled RGD have revealed that time, temperature, pH, 

presence of trace metal contaminants, and stoichiometric ratio of chelator to isotope all have 

significant effects on the rate of chelation and radiolabeling efficiency. 175 A major challenge in 

development of therapeutic radiopharmaceuticals is radiolytic degradation of radiolabeled products 

because of production of free radicals in the presence of a large amount of high energy 

b-particles. 176 A study on the stability of 90 Y labeled, dimeric RGD peptide showed that presence 

TABLE 9.5 

Conjugates for Angiogenesis Targeted Radiotherapy 


Compound Radiolabel Target Tumor Model Reference 

DOTA-E- c(RGDfK)2 Y-90 aVb3 integrin OVCAR-3 human ovarian carcinoma 107, 205 

DTPA- Tyr3-octreotate-c(RGDyD) In-111 aVb3 integrin CA20948 & AR42J rat 

pancreatic tumor 

206,207 

HPMA copolymer-RGD4C Y-90 aVb3 integrin DU145 human prostate carcinoma 108 



of free radical scavengers such as gentisic acid (GA) and ascorbic acid (AA), together with storage 

at low temperature (K788C), stabilized the labeled compound for at least two half-lives of 90 Y, 

even at a high specific activity. 177 

To develop an HPMA copolymer, RGD-based targeted angiogenesis imaging and therapy 

agent, the biodistribution and tumor accumulation of HPMA copolymer-RGD4C conjugate as 

compared to a control conjugate containing the tripeptide Arg-Gly-Glu (HPMA copolymer- 

RGE4C conjugate) has been studied. 178 Figure 9.3 shows the scintigraphic image of SCID mice 

bearing prostate tumor xenografts 24 h post-injection of the conjugates. The HPMA copolymer- 

RGD4C conjugate shows greater tumor accumulation than the control. The tumor uptake of HPMA 

copolymer-RGD4C conjugate (4.6G1.8% injected dose/g) at 24 h post-injection was significantly 

higher than the control (1.2G0.1%). A time-dependent biodistribution study showed sustained 

tumor accumulation of HPMA copolymer-RGD4C conjugate, efficient background clearance, 

and increasing T/B ratios over time. High T/B increases both tumor detection and therapeutic 

ratio. 178 Another advantage of using a polymeric conjugate of RGD was highlighted by comparing 

the biodistributions of HPMA copolymer-RGD4C conjugates and free RGD4C. 23 Organ distribution 

data in two murine xenograft human prostate models indicated higher tumor accumulation 

and lower extravasation in normal tissues for HPMA copolymer conjugates compared to free 

RGD4C peptides (Figure 9.4). Also, T/B was significantly higher for the macromolecular conjugate. 

These conjugates may be particularly advantageous for cancer radiotherapy because the 

combination of polyvalent interaction and EPR effect would help to retain the conjugate in the 

tumor, enhancing the radiation dose. 108,158 

The use of a water-soluble polymer (HPMA)-based conjugate of RGD peptide and the acyclic 

chelator cyclohexyl-diethylenetriamine pentaacetic acid (CHX-A 00 -DTPA) for angiogenesis 

directed ( 90 Y) radiotherapy has been studied. 108 After intravenous injection in prostate carcinoma 

(a) (b) 

FIGURE 9.3 (See color insert following page 522.) Scintigraphic images of human prostate tumor-bearing 

SCID mice 24 h post-intravenous injection of 99m Tc labeled HPMA copolymer conjugates. HPMA copolymer- 

RGD4C conjugate shows higher localization in tumor as compared to the control, HPMA copolymer-RGE4C 

conjugate (solid arrow). Additional higher activity was found in the kidney (broken arrow). The mouse 

radiograph (center) shows anatomic correlation of tumor and other organs. Figure legends: (a) HPMA copolymer-RGD4C 

conjugates; (b) HPMA copolymer-RGE4C conjugates. (From Mitra, A., et al., J. Control. 

Release, 102, 191, 2005. With permission.) 


172 

% ID/g 

(a) 

% ID/g 

(b) 

16.000 

14.000 

12.000 

10.000 

8.000 

6.000 

4.000 

2.000 

0.000 

16.000 

14.000 

12.000 

10.000 

8.000 

6.000 

4.000 

2.000 

0.000 

HPMA-RGD4C (n=6) 

RGD4C-DPK (n=6) 

HPMA-RGE4C (n=6) 

RGE4C-DPK (n=6) 

Blood Heart Lung Liver Spleen Kidney Muscle Tumor 

HPMA-RGD4C (n=6) 

RGD4C-DPK (n=6) 

HPMA-RGE4C (n=6) 

RGE4C-DPK (n=6) 


Blood Heart Lung Liver Spleen Kidney Muscle Tumor 

FIGURE 9.4 Residual radioactivity in % injected dose per gram of organ tissue (%ID/g) 24 h post-intravenous 

injection of 99m Tc labeled HPMA copolymer conjugates and free peptides in (a) DU145 and (b) PC-3 prostate 

tumor xenograft-bearing SCID mice. The HPMA copolymers showed significantly higher tumor accumulation 

and reduced background localization than the peptides. The organ data are expressed as meanGSD (number of 

animals/group is shown). (From Line, B. R., et al., J. Nucl. Med., 46, 1152, 2005. With permission.) 

xenograft bearing SCID mice, the tumor accumulation of the conjugate peaked at 72 h post-injection, 

whereas the accumulation in other major organs significantly decreased during that period. A single 

injection of the 90 Y labeled conjugate at dose levels of 100 and 250 Ci caused significant reduction 

of tumor volume as compared to the untreated control that was evident from day 7 post-injection 

(Figure 9.5, top panel). Tumor histopathology showed cellular apoptosis in the treated groups, 

whereas no signs of acute toxicity were observed in the kidney, liver, and spleen (Figure 9.5, bottom 

panel). 

The RGD4C peptide has a doubly cyclized structure containing two disulphide bonds that 

affords high binding affinity and specificity for the aVb3 integrin. 178,179 However, the solution 

instability of the RGD4C disulphide bonds is a potential disadvantage that can lead to significant 

reduction in aVb3 binding affinity. 180 To overcome this problem, highly stable RGD peptides 

having monocyclic structure with head-to-tail cyclization containing a D-amino acid, e.g., 

RGDfK or RGDyK, have been synthesized. 117 These have high affinity for aVb3 and have been 

used for delivery of imaging agents and therapeutics to tumor angiogenesis. 19,21 The long term (up 

to 192 h) biodistribution and tumor targeting properties of multivalent HPMA copolymer 



Tumor volume (cm 3 ) 

2.00 

1.80 

1.60 

1.40 

1.20 

1.00 

0.80 

0.60 

0.40 

0.20 

0.00 

0 2 4 6 8 10 12 14 16 18 20 22 

Days post-treatment 

(a) (b) 

(c) (d) 

FIGURE 9.5 (See color insert following page 522.) TOP PANEL: Effect of 90 Y labeled HPMA copolymer- 

RGD4C conjugate treatment on human prostate tumor (DU145) growth in SCID mice. Animal groups treated 

with single dose of 100 mCi (p!0.03) and 250 mCi (p!0.01) 90 Y-HPMA-RGD4C conjugate showed significant 

tumor growth reduction as compared to the untreated controls by day seven post-treatment. Data are 

presented as mean of tumor volume (cm 3 )GSD (nZ6 mice per group). Figure legends: untreated control 

(closed square), 250 mCi of 90 Y-HPMA copolymer-RGD4C conjugate (closed circle), and 100 mCi of 90 Y- 

HPMA copolymer-RGD4C conjugate (open triangle). BOTTOM PANEL: Histological analysis of kidney and 

tumor samples at 21 days post-treatment following injection of 250 mCi 90 Y labeled HPMA copolymer- 

RGD4C conjugate or untreated control. Kidney samples taken from (a) control group and (b) 250 mCi treatment 

group were similar and showed no radiation induced toxicity. There was a complete lack of evidence of 

any tubular epithelial injury. Normal glomerular and proximal tubular anatomy was evident in both control and 

90 Y treated specimens. (c) The tumor sections from the control animals showed high grade epithelial malignancy 

typical of DU145 xenografts. (d) Tumor samples from 250 mCi treatment animals showed large cellular 

drop out areas (black arrow), higher numbers of eosinophilic cytoplasmic hyaline globular bodies (thanatosomes, 

open arrow), and pronounced nuclear atypia (hatched arrow) indicative of treatment effect/induced cell 

damage (From Mitra, A., et al., Nucl. Med. Biol., 33, 43, 2006. With permission.) 


174 

conjugates of RGD4C and RGDfK peptides have been studied. 181 Scintigraphic images and 

necropsy organ counts (Figure 9.6) showed that tumor accumulation of both HPMA-RGD4C 

and HPMA-RGDfK conjugates increased over time with peak accumulations at 4.9G0.9% (96 h 

p.i.) and 5.0G1.2% (48 h p.i.) ID/g, respectively. In contrast, the background organ distribution 

(a) 

(b) 

(c) 

% ID/g 

% ID/g 

10.000 

9.000 

8.000 

7.000 

6.000 

5.000 

4.000 

3.000 

2.000 

1.000 

0.000 

10.000 

9.000 

8.000 

7.000 

6.000 

5.000 

4.000 

3.000 

2.000 

1.000 

0.000 

15min (n=3) 

1h (n=3) 

6h (n=3) 

24h (n=4) 

48h (n=4) 

96h (n=4) 

192h (n=3) 

Blood Heart Lung Liver Spleen Kidney Small 

lntestine 

15min (n=3) 

1h (n=3) 

6h (n=3) 

24h (n=4) 

48h (n=4) 

96h (n=4) 

192h (n=3) 

Blood Heart Lung Liver Spleen Kidney Small 

lntestine 


Large 

lntestine 

Large 

lntestine 

Stomach Tail Muscle Tumor 

Stomach Tail Muscle Tumor 

FIGURE 9.6 (See color insert following page 522.) (a) Scintigraphic images of Lewis lung carcinomabearing 

mice up to 192 h post-intravenous injection of 111 In labeled copolymers. (1) HPMA copolymer- 

RGD4C conjugate and (2) HPMA copolymer-RGDfK conjugate showed marked localization in tumor at 

24 h p.i. and thereafter (solid arrow). The mouse radiograph shows anatomic correlation of tumor and other 

organs. Biodistribution of (b) 111 In-HPMA copolymer-RGD4C conjugate and (c) 111 In-HPMA copolymer- 

RGDfK conjugate. The organ activities, expressed as % injected dose per gram of organ tissue (%ID/g), 

showed persistent tumor localization and clearance of activity from the background organs (From Mitra, 

A., et al., J. Control. Release, 114, 175, 2006. With permission.) 



rapidly cleared over time, resulting in significant increase in T/B ratios. The radioactive dose to 

organs as indicated by the area under curve was highest for the tumor. The polymer conjugates of 

RGD4C or RGDfK provides a means to enhance tumor uptake, decrease background accumulation, 

and enable selective delivery of therapeutic or diagnostic agents to tumor sites. Because both 

HPMA-RGD4C and HPMA-RGDfK conjugates have similar tumor targeting abilities and pharmacokinetics, 

RGDfK can be a suitable ligand because of higher solution stability and easier synthetic 

manipulation than the 12 amino-acid RGD4C. 

These studies with HPMA copolymer-RGD conjugates demonstrate the feasibility and advantages 

of using multivalent polymer-peptide conjugates. These conjugates show prolonged retention 

at the tumor site and enhanced T/B ratios. This increased contrast at the tumor site will be necessary 

for development of a clinically relevant diagnostic agent for therapy planning. Further, the 

improved therapeutic index will be ideal for angiogenesis directed chemo- or radiotherapy. 

9.5 SUMMARY AND FUTURE DIRECTIONS 

There is a tremendous interest in targeted drug delivery to tumor vasculature given the genetic 

stability and accessibility of angiogenesis markers and the importance of angiogenesis in tumor 

growth and metastasis. A number of ligands targeting angiogenic markers have been identified, but 

the major focus has been on targeting aVb3 integrins using RGD peptides. 21 Many reports of 

conjugating RGD ligands to a diverse group of macromolecular carrier systems have been 

published. These include water-soluble polymers (e.g., PEG, 139 N-(2-hydroxypropyl) methacrylamide, 

178 and polyethylenimine 164 ), proteins, 182,183 liposomes, 184,185 and nanoparticles. 165,186 

With respect to future development, a broad range of angiogenesis targeting polymer conjugates 

can be envisioned. As additional vascular targets are identified by in vitro and in vivo 

methods, angiogenesis targeting may become a critically important strategy for fighting neoplastic 

as well as non-neoplastic diseases. The first successful human trials for imaging aVb3 integrins 

using [ 18 F]-galacto-RGD 124,125 and the results of pre-clinical studies using macromolecular conjugates 

of RGD argue the validity of this concept. 

It is expected that the biokinetics of copolymer-conjugates will be continuously improved by 

tailoring the molecular weight 187–190 and/or electronegative charge 190,191 of the conjugates. This 

flexibility in the design and synthesis of HPMA and other copolymer conjugates is likely to be 

important in modifying the blood half-life or reducing non-specific accumulation in normal organs, 

factors that should increase the conjugate therapeutic index. These may incorporate either enzymatically 

hydrolysable (e.g., tetrapeptide GFLG) 143 or pH sensitive (e.g., hydrazone linker) 192 

linkers for intracellular drug release. In addition, they will likely demonstrate targeting ligand 

multivalency to increase the tumor uptake and therapeutic efficacy. 

It is also expected that there will be expanded use of therapeutic radionuclides to destroy 

angiogenic vasculature and surrounding tumor cells. Because radioactive emissions can kill at a 

distance from the point of radioisotope localization, they have a diameter of effectiveness that may 

overcome the problem of tumor heterogeneity that has plagued other molecular therapies. In short, 

the ever increasing interest in polymer-based therapeutics promises continued progress in angiogenesis 

targeted, tumor imaging, and therapy. 

9.6 LIST OF ABBREVIATIONS 

AFK Ala-Phe-Lys 

APA Aminopeptidase A 

APMA-CHX-A 00 -DTPA N-methacryloylaminopropyl-2-amino-3-(isothiourea-phenyl)propyl-cyclohexane-1,2-diamine-N,N-N 

0 ,N 0 ,N 00 ,N 00 - 

pentaacetic acid 


176 


APN Aminopeptidase N 

BBB Blood–brain barrier 

CAT Chloramphenicol acetyl transferase 

CDTRL Cys-Asp-Thr-Arg-Leu 

CGKRK Cys-Gly-Lys-Arg-Lys 

CHX-A 00 -DTPA Cyclohexyl–diethylenetriamine pentaacetic acid 

CPRECESIC Cys-Pro-Arg-Glu-Cys-Glu-Ser-Ile-Cys 

CRGRRST Cys-Arg-Gly-Arg-Arg-Ser-Thr 

CSRPRRSEC Cys-Ser-Arg-Pro-Arg-Arg-Ser-Glu-Cys 

DOTA 1,4,7,10-tetra-azacylcododecane-N,N 0 ,N 00 ,N 000 -tetraacetic acid 

DOX Doxorubicin 

DPK N-u-bis(2-pyridylmethyl)-L-lysine 

DTPA Diethylenetriaminepentaacetic acid 

ECM Extracellular matrix 

EPR Enhanced permeability and retention 

FB Fluorobenzoyl 

FDA Food and drug administration 

FDG 2-Fluoro-2-deoxyglucose 

5-Fu 5-Fluorouracil 

GFLG Gly-Phe-Leu-Gly 

GPLGV Gly-Pro-Leu-Gly-Val 

GPLGIAGQ Gly-Pro-Leu-Gly-Ile-Ala-Gly-Gln 

GPLGVRGK Gly-Pro-Leu-Gly-Val-Arg-Gly-Lys 

GVPLSLTMGC Gly-Val-Pro-Leu-Ser-Leu-Thr-Met-Gly-Cys 

HPMA N-(2-hydroxypropyl) methacrylamide 

HWGF His-Trp-Gly-Phe 

HYNIC Hydrazinonicotinamide 

MA-GG-DPK N-methacryloylglycylglycyl-(N-u-bis(2-pyridylmethyl)-L-lysine) 

MA-GG-RGD4C N-methacryloylglycylglycyl-RGD4C 

MA-Tyr N-methacryloyltyrosinamide 

MBq Megabecquerel 

mCi Millicurie 

MMP Matrix metalloproteinases 

MRI Magnetic resonance imaging 

NGR Asn-Gly-Arg 

NIR Near-infrared 

PEG Polyethylene glycol 

PEI Polyethylenimine 

PET Positron emission tomography 

PLGL Pro-Leu-Gly-Leu 

PLL Poly-L-lysine 

RGD Arg-Gly-Asp 

RGE Arg-Gly-Glu 

SCID Severe combined immunodeficient 

SMSIARL Ser-Met-Ser-Ile-Ala-Arg-Leu 

T/B Tumor-to-background 

TIMP Tissue inhibitors of matrix metalloproteinases 

VEGF Vascular endothelial growth factor 



ACKNOWLEDGMENT 

Support from the American Russian Cancer Alliance and National Institutes of Health (1 R01 

EB0207171) is acknowledged. 

REFERENCES 

1. Ferrara, N. and Kerbel, R. S., Angiogenesis as a therapeutic target, Nature, 438, 967, 2005. 

2. Carmeliet, P., Angiogenesis in health and disease, Nat. Med., 9, 653, 2003. 

3. Carmeliet, P. and Jain, R. K., Angiogenesis in cancer and other diseases, Nature, 407, 249, 2000. 

4. Carmeliet, P., Mechanisms of angiogenesis and arteriogenesis, Nat. Med., 6, 389, 2000. 

5. Thorpe, P. E., Vascular targeting agents as cancer therapeutics, Clin. Cancer Res., 10, 415, 2004. 

6. Siemann, D. W. et al., Differentiation and definition of vascular-targeted therapies, Clin. Cancer 

Res., 11, 416, 2005. 

7. Kerbel, R. and Folkman, J., Clinical translation of angiogenesis inhibitors, Nat. Rev. Cancer, 2, 727, 

2002. 

8. Ellis, L. M., Synopsis of angiogenesis inhibitors in oncology, Oncology (Huntingt), 16, 14, 2002. 

9. Risau, W., Mechanisms of angiogenesis, Nature, 386, 671, 1997. 

10. Gasparini, G. et al., Angiogenic inhibitors: A new therapeutic strategy in oncology, Nat. Clin. Pract. 

Oncol., 2, 562, 2005. 

11. Folkman, J., Angiogenesis inhibitors: A new class of drugs, Cancer Biol. Ther., 2, S127, 2003. 

12. Ranieri, G. and Gasparini, G., Angiogenesis and angiogenesis inhibitors: A new potential anticancer 

therapeutic strategy, Curr. Drug Targets Immune Endocr. Metabol. Disord., 1, 241, 2001. 

13. Ferrara, N. et al., Discovery and development of bevacizumab, an anti-VEGF antibody for treating 

cancer, Nat. Rev. Drug Discov., 3, 391, 2004. 

14. Tozer, G. M., Kanthou, C., and Baguley, B. C., Disrupting tumour blood vessels, Nat. Rev. Cancer, 

5, 423, 2005. 

15. Narazaki, M. and Tosato, G., Targeting coagulation to the tumor microvasculature: Perspectives and 

therapeutic implications from preclinical studies, J. Natl. Cancer Inst., 97, 705, 2005. 

16. Ching, L. M. et al., Induction of intratumoral tumor necrosis factor (TNF) synthesis and hemorrhagic 

necrosis by 5,6-dimethylxanthenone-4-acetic acid (DMXAA) in TNF knockout mice, Cancer Res., 

59, 3304, 1999. 

17. Tozer, G. M. et al., The biology of the combretastatins as tumour vascular targeting agents, Int. 

J. Exp. Pathol., 83, 21, 2002. 

18. Pedley, R. B. et al., Eradication of colorectal xenografts by combined radioimmunotherapy and 

combretastatin a-4 3-O-phosphate, Cancer Res., 61, 4716, 2001. 

19. Haubner, R. and Wester, H. J., Radiolabeled tracers for imaging of tumor angiogenesis and evaluation 

of anti-angiogenic therapies, Curr. Pharm. Des., 10, 1439, 2004. 

20. Li, W. P. and Anderson, C. J., Imaging matrix metalloproteinase expression in tumors, Q. J. Nucl. 

Med., 47, 201, 2003. 

21. Temming, K. et al., RGD-based strategies for selective delivery of therapeutics and imaging agents 

to the tumour vasculature, Drug Resist.Updat., 8 (6), 381–402, 2005. 


review, J. Control. Release, 65, 271, 2000. 

23. Line, B. R. et al., Targeting tumor angiogenesis: Comparison of peptide and polymer-peptide 

conjugates, J. Nucl. Med., 46, 1552, 2005. 

24. Mahadevan, V. and Hart, I. R., Metastasis and angiogenesis, Acta Oncol., 29, 97, 1990. 

25. Hanahan, D. and Folkman, J., Patterns and emerging mechanisms of the angiogenic switch during 

tumorigenesis, Cell, 86, 353, 1996. 

26. Folkman, J., Tumor angiogenesis: Therapeutic implications, N. Engl. J. Med., 285, 1182, 1971. 

27. Folkman, J., What is the evidence that tumors are angiogenesis dependent?, J. Natl. Cancer Inst., 82, 

4, 1990. 

28. Weidner, N., Tumoural vascularity as a prognostic factor in cancer patients: The evidence continues 

to grow, J. Pathol., 184, 119, 1998. 


178 


29. Weidner, N., Tumour vascularity and proliferation: Clear evidence of a close relationship, J. Pathol., 

189, 297, 1999. 

30. Bergers, G. and Benjamin, L. E., Tumorigenesis and the angiogenic switch, Nat. Rev. Cancer, 3, 401, 

2003. 

31. Hanahan, D. and Weinberg, R. A., The hallmarks of cancer, Cell, 100, 57, 2000. 

32. Folkman, J., Endogenous angiogenesis inhibitors, APMIS, 112, 496, 2004. 

33. Nyberg, P., Xie, L., and Kalluri, R., Endogenous inhibitors of angiogenesis, Cancer Res., 65, 3967, 

2005. 

34. Sund, M. et al., Function of endogenous inhibitors of angiogenesis as endothelium-specific tumor 

suppressors, Proc. Natl. Acad. Sci. U.S.A., 102, 2934, 2005. 

35. Relf, M. et al., Expression of the angiogenic factors vascular endothelial cell growth factor, acidic 

and basic fibroblast growth factor, tumor growth factor beta-1, platelet-derived endothelial cell 

growth factor, placenta growth factor, and pleiotrophin in human primary breast cancer and its 

relation to angiogenesis, Cancer Res., 57, 963, 1997. 

36. Yancopoulos, G. D. et al., Vascular-specific growth factors and blood vessel formation, Nature, 407, 

242, 2000. 

37. Hanahan, D. et al., Transgenic mouse models of tumour angiogenesis: The angiogenic switch, its 

molecular controls, and prospects for preclinical therapeutic models, Eur. J Cancer, 32A, 2386, 

1996. 

38. Harris, A. L., Hypoxia—a key regulatory factor in tumour growth, Nat. Rev. Cancer, 2, 38, 2002. 

39. Semenza, G. L., HIF-1 and tumor progression: Pathophysiology and therapeutics, Trends Mol. Med., 

8, S62, 2002. 

40. Jain, R. K., Munn, L. L., and Fukumura, D., Dissecting tumour pathophysiology using intravital 

microscopy, Nat. Rev. Cancer, 2, 266, 2002. 

41. Morikawa, S. et al., Abnormalities in pericytes on blood vessels and endothelial sprouts in tumors, 

Am. J. Pathol., 160, 985, 2002. 

42. Jain, R. K., Molecular regulation of vessel maturation, Nat. Med., 9, 685, 2003. 

43. Baluk, P., Hashizume, H., and McDonald, D. M., Cellular abnormalities of blood vessels as targets in 

cancer, Curr. Opin. Genet. Dev., 15, 102, 2005. 

44. Ruoslahti, E., Specialization of tumour vasculature, Nat. Rev. Cancer, 2, 83, 2002. 

45. Brekken, R. A. and Thorpe, P. E., Vascular endothelial growth factor and vascular targeting of solid 

tumors, Anticancer Res., 21, 4221, 2001. 

46. Burrows, F. J. and Thorpe, P. E., Vascular targeting—a new approach to the therapy of solid tumors, 

Pharmacol. Ther., 64, 155, 1994. 

47. Neri, D. and Bicknell, R., Tumour vascular targeting, Nat. Rev. Cancer, 5, 436, 2005. 


Nature, 380, 364, 1996. 

49. Arap, W. et al., Steps toward mapping the human vasculature by phage display, Nat. Med., 8, 121, 

2002. 

50. Ruoslahti, E., Vascular zip codes in angiogenesis and metastasis, Biochem. Soc. Trans., 32, 397, 

2004. 

51. Haubner, R. H. et al., Radiotracer-based strategies to image angiogenesis, Q. J. Nucl. Med., 47, 189, 

2003. 

52. McQuade, P. and Knight, L. C., Radiopharmaceuticals for targeting the angiogenesis marker 

alpha(v)beta(3), Q. J. Nucl. Med., 47, 209, 2003. 

53. Pasqualini, R., Koivunen, E., and Ruoslahti, E., Alpha v integrins as receptors for tumor targeting by 

circulating ligands, Nat. Biotechnol., 15, 542, 1997. 

54. Pasqualini, R. et al., Aminopeptidase N is a receptor for tumor-homing peptides and a target for 

inhibiting angiogenesis, Cancer Res., 60, 722, 2000. 

55. Isberg, R. R. and Tran, V. N., Binding and internalization of microorganisms by integrin receptors, 

Trends Microbiol., 2, 10, 1994. 

56. Hart, S. L. et al., Cell binding and internalization by filamentous phage displaying a cyclic Arg-Gly- 

Asp-containing peptide, J. Biol. Chem., 269, 12468, 1994. 

57. Ferrara, N., Gerber, H. P., and LeCouter, J., The biology of VEGF and its receptors, Nat. Med., 9, 

669, 2003. 



58. Ferrara, N., Vascular endothelial growth factor: Basic science and clinical progress, Endocr. Rev., 

25, 581, 2004. 

59. Ferrara, N., Vascular endothelial growth factor as a target for anticancer therapy, Oncologist, 9 

(Suppl 1), 2, 2004. 

60. Ahmed, S. I., Thomas, A. L., and Steward, W. P., Vascular endothelial growth factor (VEGF) 

inhibition by small molecules, J. Chemother., 16 (Suppl 4), 59, 2004. 

61. Sato, Y., Aminopeptidases and angiogenesis, Endothelium, 10, 287, 2003. 

62. Chang, Y. W. et al., CD13 (aminopeptidase N) can associate with tumor-associated antigen L6 and 

enhance the motility of human lung cancer cells, Int. J. Cancer, 116, 243, 2005. 

63. Ishii, K. et al., Aminopeptidase N regulated by zinc in human prostate participates in tumor cell 

invasion, Int. J. Cancer, 92, 49, 2001. 

64. Arap, W., Pasqualini, R., and Ruoslahti, E., Cancer treatment by targeted drug delivery to tumor 

vasculature in a mouse model, Science, 279, 377, 1998. 

65. Ellerby, H. M. et al., Anti-cancer activity of targeted pro-apoptotic peptides, Nat. Med., 5, 1032, 

1999. 

66. Schlingemann, R. O. et al., Aminopeptidase a is a constituent of activated pericytes in angiogenesis, 

J. Pathol., 179, 436, 1996. 

67. Marchio, S. et al., Aminopeptidase A is a functional target in angiogenic blood vessels, Cancer Cell, 

5, 151, 2004. 

68. Vihinen, P. and Kahari, V. M., Matrix metalloproteinases in cancer: Prognostic markers and therapeutic 

targets, Int. J. Cancer, 99, 157, 2002. 

69. Vihinen, P., Ala-aho, R., and Kahari, V. M., Matrix metalloproteinases as therapeutic targets in 

cancer, Curr. Cancer Drug Targets, 5, 203, 2005. 

70. Egeblad, M. and Werb, Z., New functions for the matrix metalloproteinases in cancer progression, 

Nat. Rev. Cancer, 2, 161, 2002. 

71. Nguyen, M., Arkell, J., and Jackson, C. J., Human endothelial gelatinases and angiogenesis, Int. 

J. Biochem. Cell Biol., 33, 960, 2001. 

72. Himelstein, B. P. et al., Metalloproteinases in tumor progression: The contribution of MMP-9, Invas. 

Metast., 14, 246, 1994. 

73. Koivunen, E. et al., Tumor targeting with a selective gelatinase inhibitor, Nat. Biotechnol., 17, 768, 

1999. 

74. Hidalgo, M. and Eckhardt, S. G., Development of matrix metalloproteinase inhibitors in cancer 

therapy, J. Natl. Cancer Inst., 93, 178, 2001. 

75. Levitt, N. C. et al., Phase I and pharmacological study of the oral matrix metalloproteinase inhibitor, 

MMI270 (CGS27023A), in patients with advanced solid cancer, Clin. Cancer Res., 7, 1912, 2001. 

76. Eatock, M. et al., A dose-finding and pharmacokinetic study of the matrix metalloproteinase 

inhibitor MMI270 (previously termed CGS27023A) with 5-FU and folinic acid, Cancer Chemother. 

Pharmacol., 55, 39, 2005. 

77. Kline, T., Torgov, M. Y., Mendelsohn, B. A., Cerveny, C. G., and Senter, P. D., Novel antitumor 

prodrugs designed for activation by matrix metalloproteinases-2 and -9, Mol. Pharm., 1, 9, 2004. 

78. Albright, C. F. et al., Matrix metalloproteinase-activated doxorubicin prodrugs inhibit HT1080 

xenograft growth better than doxorubicin with less toxicity, Mol. Cancer Ther., 4, 751, 2005. 

79. Bae, M. et al., Metalloprotease-specific poly(ethylene glycol) methyl ether–peptide–doxorubicin 

conjugate for targeting anticancer drug delivery based on angiogenesis, Drugs Exp. Clin. Res., 29, 

15, 2003. 

80. Kratz, F. et al., Development and in vitro efficacy of novel MMP2 and MMP9 specific doxorubicin 

albumin conjugates, Bioorg. Med. Chem. Lett., 11, 2001, 2001. 

81. Kratz, F. et al., Probing the cysteine-34 position of endogenous serum albumin with thiol-binding 

doxorubicin derivatives. Improved efficacy of an acid-sensitive doxorubicin derivative with specific 

albumin-binding properties compared to that of the parent compound, J. Med. Chem., 45, 5523, 

2002. 

82. Mansour, A. M. et al., A new approach for the treatment of malignant melanoma: Enhanced antitumor 

efficacy of an albumin-binding doxorubicin prodrug that is cleaved by matrix 

metalloproteinase 2, Cancer Res., 63, 4062, 2003. 


180 


83. Kuhnast, B. et al., Targeting of gelatinase activity with a radiolabeled cyclic HWGF peptide, Nucl. 

Med. Biol., 31, 337, 2004. 

84. Zheng, Q. H. et al., Synthesis and preliminary biological evaluation of MMP inhibitor radiotracers 

[ 11 C]methyl-halo-CGS 27023A analogs, new potential PET breast cancer imaging agents, Nucl. 

Med. Biol., 29, 761, 2002. 

85. Furumoto, S., Iwata, R., and Ido, T., Design and synthesis of fluorine-18 labeled matrix metalloproteinase 

inhibitors for cancer imaging, J. Labelled Compd. Rad., 45, 975, 2002. 

86. Fei, X., Hutchins, G. D. et al., Synthesis of radiolabeled biphenylsulfonamide matrix metalloproteinase 

inhibitors as new potential PET cancer imaging agents, Bioorg. Med. Chem. Lett., 13, 2217, 

2003. 

87. Furumoto, S. et al., Development of a new 18 F-labeled matrix metalloproteinase-2 inhibitor for 

cancer imaging by PET, J. Nucl. Med., 43 (Suppl), 364, 2002. 

88. Furumoto, S. et al., Tumor detection using 18 F-labeled matrix metalloproteinase-2 inhibitor, Nucl. 

Med. Biol., 30, 119, 2003. 

89. Zheng, Q. H. et al., Synthesis, biodistribution and micro-PET imaging of a potential cancer biomarker 

carbon-11 labeled MMP inhibitor ( 2 R)-2-[[4-(6-fluorohex-1-ynyl)phenyl]sulfonylamino]-3methylbutyric 

acid [ 11 C]methyl ester, Nucl. Med. Biol., 30, 753, 2003. 

90. Li, W. P. et al., 64 Cu-DOTA-CTTHWGFTLC: A selective gelatinase inhibitor for tumor imaging, 

J. Nucl. Med., 43 (Suppl), 227, 2002. 

91. Li, W. P. et al., In vitro and in vivo evaluation of a radiolabeled gelatinase inhibitor for microPET 

imaging of metastatic breast cancer, Mol. Imaging Biol., 4 (Suppl), 23, 2002. 

92. Giersing, B. K. et al., Synthesis and characterization of 111 In-DTPA-N-TIMP-2: A radiopharmaceutical 

for imaging matrix metalloproteinase expression, Bioconjug. Chem., 12, 964, 2001. 

93. Kulasegaram, R. et al., In vivo evaluation of 111 In-DTPA-N-TIMP-2 in Kaposi sarcoma associated 

with HIV infection, Eur. J. Nucl. Med., 28, 756, 2001. 

94. Bremer, C., Tung, C. H., and Weissleder, R., In vivo molecular target assessment of matrix metalloproteinase 

inhibition, Nat. Med., 7, 743, 2001. 

95. Bremer, C. et al., Optical imaging of matrix metalloproteinase-2 activity in tumors: Feasibility study 

in a mouse model, Radiology, 221, 523, 2001. 

96. Bremer, C., Tung, C. H., and Weissleder, R., Molecular imaging of MMP expression and therapeutic 

MMP inhibition, Acad. Radiol., 9 (Suppl 2), S314, 2002. 

97. Pham, W. et al., Developing a peptide-based near-infrared molecular probe for protease sensing, 

Bioconjug. Chem., 15, 1403, 2004. 

98. Cox, D. et al., The pharmacology of the integrins, Med. Res. Rev., 14, 195, 1994. 

99. van Der, F. A. and Sonnenberg, A., Function and interactions of integrins, Cell Tissue Res., 305, 285, 

2001. 

100. Sepulveda, J. L., Gkretsi, V., and Wu, C., Assembly and signaling of adhesion complexes, Curr. Top. 

Dev. Biol., 68, 183, 2005. 

101. Hood, J. D. and Cheresh, D. A., Role of integrins in cell invasion and migration, Nat. Rev. Cancer,2, 

91, 2002. 

102. Giancotti, F. G. and Ruoslahti, E., Integrin signaling, Science, 285, 1028, 1999. 

103. Brooks, P. C., Clark, R. A., and Cheresh, D. A., Requirement of vascular integrin alpha v beta 3 for 

angiogenesis, Science, 264, 569, 1994. 

104. Friedlander, M. et al., Definition of two angiogenic pathways by distinct alpha v integrins, Science, 

270, 1500, 1995. 

105. Eliceiri, B. P. and Cheresh, D. A., Role of alpha v integrins during angiogenesis, Cancer J., 6 (Suppl 

3), S245, 2000. 

106. Stupack, D. G. and Cheresh, D. A., Integrins and angiogenesis, Curr. Top. Dev. Biol., 64, 207, 2004. 

107. Janssen, M. L. et al., Tumor targeting with radiolabeled alpha(v)beta(3) integrin binding peptides in a 

nude mouse model, Cancer Res., 62, 6146, 2002. 

108. Mitra, A. et al., Polymer-peptide conjugates for angiogenesis targeted tumor radiotherapy, Nucl. 

Med. Biol., 33, 43, 2006. 

109. Goodman, S. L. et al., Nanomolar small molecule inhibitors for alphav(beta)6, alphav(beta)5, and 

alphav(beta)3 integrins, J. Med. Chem., 45, 1045, 2002. 



110. Winter, P. M. et al., Molecular imaging of angiogenesis in nascent Vx-2 rabbit tumors using a novel 

alpha(nu)beta3-targeted nanoparticle and 1.5 tesla magnetic resonance imaging, Cancer Res., 63, 

5838, 2003. 

111. Chen, X., Conti, P. S., and Moats, R. A., In vivo near-infrared fluorescence imaging of integrin 

alphavbeta3 in brain tumor xenografts, Cancer Res., 64, 8009, 2004. 

112. Haubner, R., Finsinger, D., and Kessler, H., Stereoisomeric peptide libraries and peptidomimetics for 

designing selective inhibitors of the alphav beta3 integrin for a new cancer therapy, Angew. Chem. 

Int. Ed. Engl., 36, 1374, 1997. 

113. Aumailley, M. et al., Arg-Gly-Asp constrained within cyclic pentapeptides. Strong and selective 

inhibitors of cell adhesion to vitronectin and laminin fragment P1, FEBS Lett., 291, 50, 1991. 

114. Gurrath, M. et al., Conformation/activity studies of rationally designed potent anti-adhesive RGD 

peptides, Eur. J. Biochem., 210, 911, 1992. 

115. Haubner, R. et al., Structural and functional aspects of RGD containing cyclic pentapeptides as 

highly potent and selective avb3 antagonists, J. Am. Chem. Soc., 118, 7461, 1996. 

116. Haubner, R. et al., Comparison of tumor uptake and biokinetics of I-125 and F-18 labeled RGDpeptides, 

J. Labelled Compd. Rad., 42, S36, 1999. 

117. Haubner, R. et al., Radiolabeled alpha(v)beta3 integrin antagonists: A new class of tracers for tumor 

targeting, J. Nucl. Med., 40, 1061, 1999. 

118. Chen, X. et al., MicroPET and autoradiographic imaging of breast cancer alpha v-integrin expression 

using 18 F- and 64 Cu-labeled RGD peptide, Bioconjug. Chem., 15, 41, 2004. 

119. Wang, W. et al., Convenient solid-phase synthesis of diethylenetriaminepenta–acetic acid (DTPA)conjugated 

cyclic RGD peptide analogues, Cancer Biother. Radiopharm., 20, 547, 2005. 

120. Haubner, R. et al., Glycosylated RGD-containing peptides: Tracer for tumor targeting and angiogenesis 

imaging with improved biokinetics, J. Nucl. Med., 42, 326, 2001. 

121. Haubner, R. et al., Noninvasive imaging of alpha(v)beta3 integrin expression using 18 F-labeled 

RGD-containing glycopeptide and positron emission tomography, Cancer Res., 61, 1781, 2001. 

122. Haubner, R. et al., [ 18 F]Galacto-RGD: Synthesis, radiolabeling, metabolic stability, and radiation 

dose estimates, Bioconjug. Chem., 15, 61, 2004. 

123. Haubner, R. et al., [F-18]-RGD-peptides conjugated with hydrophilic tetrapeptides for the noninvasive 

determination of the avb3 integrin, J. Nucl. Med., 43 (Suppl), 89P, 2002. 

124. Haubner, R. et al., Noninvasive visualization of the activated alphavbeta3 integrin in cancer patients 

by positron emission tomography and [(18)F]Galacto-RGD, PLoS Med., 2, e70, 2005. 

125. Beer, A. J. et al., Biodistribution and pharmacokinetics of the alphavbeta3-selective tracer 18 Fgalacto-RGD 

in cancer patients, J. Nucl. Med., 46, 1333, 2005. 

126. Sipkins, D. A. et al., Detection of tumor angiogenesis in vivo by alphaVbeta3-targeted magnetic 

resonance imaging, Nat. Med., 4, 623, 1998. 

127. Anderson, S. A. et al., Magnetic resonance contrast enhancement of neovasculature with alpha(v)beta(3)-targeted 

nanoparticles, Magn. Reson. Med., 44, 433, 2000. 

128. Winter, P. M. et al., Molecular imaging of angiogenesis in early-stage atherosclerosis with alpha(v)beta3-integrin-targeted 

nanoparticles, Circulation, 108, 2270, 2003. 

129. Cheng, Z. et al., Near-infrared fluorescent RGD peptides for optical imaging of integrin alphavbeta3 

expression in living mice, Bioconjug. Chem., 16, 1433, 2005. 

130. de Groot, F. M. et al., Design, synthesis, and biological evaluation of a dual tumor-specific motive 

containing integrin-targeted plasmin-cleavable doxorubicin prodrug, Mol. Cancer Ther., 1, 901, 

2002. 

131. Devy, L. et al., Plasmin-activated doxorubicin prodrugs containing a spacer reduce tumor growth 

and angiogenesis without systemic toxicity, FASEB J., 18, 565, 2004. 

132. Kim, J. W. and Lee, H. S., Tumor targeting by doxorubicin-RGD-4C peptide conjugate in an 

orthotopic mouse hepatoma model, Int. J. Mol. Med., 14, 529, 2004. 

133. Janssen, M. et al., Comparison of a monomeric and dimeric radiolabeled RGD-peptide for tumor 

targeting, Cancer Biother. Radiopharm., 17, 641, 2002. 

134. Liu, S. et al., 99m Tc-labeling of a hydrazinonicotinamide-conjugated vitronectin receptor antagonist 

useful for imaging tumors, Bioconjug. Chem., 12, 624, 2001. 

135. Chen, X. et al., MicroPET imaging of breast cancer alphav-integrin expression with 64 Cu-labeled 

dimeric RGD peptides, Mol. Imaging Biol., 6, 350, 2004. 


182 


136. Chen, X. et al., Micro-PET imaging of alphavbeta3-integrin expression with 18 F-labeled dimeric 

RGD peptide, Mol. Imaging, 3, 96, 2004. 

137. Wu, Y. et al., MicroPET imaging of glioma integrin {alpha}v{beta}3 expression using (64)Culabeled 

tetrameric RGD peptide, J. Nucl. Med., 46, 1707, 2005. 

138. Poethko, T. et al., Two-step methodology for high-yield routine radiohalogenation of peptides: 

(18)F-labeled RGD and octreotide analogs, J. Nucl. Med., 45, 892, 2004. 

139. Chen, X. et al., Pharmacokinetics and tumor retention of 125 I-labeled RGD peptide are improved by 

PEGylation, Nucl. Med. Biol., 31, 11, 2004. 

140. Chen, X. et al., MicroPET imaging of brain tumor angiogenesis with 18 F-labeled PEGylated RGD 

peptide, Eur. J. Nucl. Med. Mol. Imaging, 31, 1081, 2004. 

141. Chen, X. et al., Pegylated Arg-Gly-Asp peptide: 64 Cu labeling and PET imaging of brain tumor 

alphavbeta3-integrin expression, J. Nucl. Med., 45, 1776, 2004. 

142. Poethko, T. et al., Improved tumor uptake, tumor retention and tumor/background ratios of pegylated 

RGD-multimers, J. Nucl. Med., 44, 46P, 2003. 

143. Putnam, D. and Kopecek, J., Polymer conjugates with anticancer activity, Adv. Polym. Sci., 122, 56, 

1995. 

144. Kopecek, J. et al., HPMA copolymer-anticancer drug conjugates: Design, activity, and mechanism 

of action, Eur. J. Pharm. Biopharm., 50, 61, 2000. 


146. Greenwald, R. B. et al., Effective drug delivery by PEGylated drug conjugates, Adv. Drug Deliv. 

Rev., 55, 217, 2003. 

147. Kopecek, J. et al., Water soluble polymers in tumor targeted delivery, J. Control. Release, 74, 147, 

2001. 

148. Lu, Z. R. et al., Design of novel bioconjugates for targeted drug delivery, J. Control. Release, 78, 

165, 2002. 

149. Duncan, R. et al., Fate of N-(2-hydroxypropyl)methacrylamide copolymers with pendent galactosamine 

residues after intravenous administration to rats, Biochim. Biophys. Acta, 880, 62, 1986. 

150. Kopecek, J. and Duncan, R., Poly[N-(2-Hydroxypropyl)-methacrylamide] macromolecules as drug 

carrier systems, In Polymers in Controlled Drug Delivery, Illum, L. and Davis, S. S. Eds., Butterworth-Heinemann, 

Oxford, UK, 1987. 

151. Lu, Z. R. et al., Functionalized semitelechelic poly[N-(2-hydroxypropyl)methacrylamide] for protein 

modification, Bioconjug. Chem., 9, 793, 1998. 

152. Mitra, A., et al., Nanocarriers for nuclear imaging and radiotherapy of cancer, Curr. Pharm. Des.,in 

press, 2006. 


chemotherapy: Mechanism of tumoritropic accumulation of proteins and the antitumor agent 


154. Satchi-Fainaro, R. et al., Targeting angiogenesis with a conjugate of HPMA copolymer and TNP- 

470, Nat. Med., 10, 255, 2004. 

155. Satchi-Fainaro, R. et al., Inhibition of vessel permeability by TNP-470 and its polymer conjugate, 

caplostatin, Cancer Cell, 7, 251, 2005. 

156. Shiah, J. J. et al., Biodistribution of free and N-(2-hydroxypropyl)methacrylamide copolymer-bound 

mesochlorin e(6) and adriamycin in nude mice bearing human ovarian carcinoma OVCAR-3 xenografts, 


157. Mammen, M. et al., Polyvalent interactions in biological systems: Implications for design and use of 

multivalent ligands and inhibitors, Angew. Chem. Int. Ed., 37, 2754, 1998. 

158. Wester, H. J. and Kessler, H., Molecular targeting with peptides or Peptide-polymer conjugates: Just 

a question of size?, J. Nucl. Med., 46, 1940, 2005. 

159. Bhargava, P. et al., A Phase I and pharmacokinetic study of TNP-470 administered weekly to 

patients with advanced cancer, Clin. Cancer Res., 5, 1989, 1999. 

160. Dickson, P. V., Nathwani, A. C., and Davidoff, A. M., Delivery of antiangiogenic agents for cancer 

gene therapy, Technol. Cancer Res. Treat., 4, 331, 2005. 

161. Tandle, A., Blazer, D. G. III, and Libutti, S. K., Antiangiogenic gene therapy of cancer: Recent 

developments, J. Transl. Med., 2, 22, 2004. 



162. Witlox, A. M. et al., Conditionally replicative adenovirus with tropism expanded towards integrins 

inhibits osteosarcoma tumor growth in vitro and in vivo, Clin. Cancer Res., 10, 61, 2004. 

163. Okada, Y. et al., Optimization of antitumor efficacy and safety of in vivo cytokine gene therapy using 

RGD fiber-mutant adenovirus vector for preexisting murine melanoma, Biochim. Biophys. Acta, 

1670, 172, 2004. 

164. Kim, W. J. et al., Soluble Flt-1 gene delivery using PEI-g-PEG-RGD conjugate for anti-angiogenesis, 


165. Hood, J. D. et al., Tumor regression by targeted gene delivery to the neovasculature, Science, 296, 

2404, 2002. 

166. Leng, Q. et al., Highly branched HK peptides are effective carriers of siRNA, J. Gene Med., 7, 977, 

2005. 

167. Pack, D. W. et al., Design and development of polymers for gene delivery, Nat. Rev. Drug Discov.,4, 

581, 2005. 

168. Kunath, K. et al., Integrin targeting using RGD-PEI conjugates for in vitro gene transfer, J. Gene 

Med., 5, 588, 2003. 

169. Erbacher, P., Remy, J. S., and Behr, J. P., Gene transfer with synthetic virus-like particles via the 

integrin-mediated endocytosis pathway, Gene Ther., 6, 138, 1999. 

170. Schiffelers, R. M. et al., Cancer siRNA therapy by tumor selective delivery with ligand-targeted 

sterically stabilized nanoparticles, Nucl. Acids Res., 32, e149, 2004. 

171. Suh, W. et al., An angiogenic, endothelial-cell-targeted polymeric gene carrier, Mol. Ther., 6, 664, 

2002. 

172. Woodle, M. C. et al., Sterically stabilized polyplex: Ligand-mediated activity, J. Control. Release, 

74, 309, 2001. 

173. Hosseinkhani, H. and Tabata, Y., PEGylation enhances tumor targeting of plasmid DNA by an 

artificial cationized protein with repeated RGD sequences, Pronectin, J. Control. Release, 97, 

157, 2004. 

174. Flynn, A. A. et al., A model-based approach for the optimization of radioimmunotherapy through 

antibody design and radionuclide selection, Cancer, 94, 1249, 2002. 

175. Liu, S. et al., (90)Y and (177)Lu labeling of a DOTA-conjugated vitronectin receptor antagonist 

useful for tumor therapy, Bioconjug. Chem., 12, 559, 2001. 

176. Garrison, W. M., Reaction mechanisms in radiolysis of peptides, polypeptides and proteins, Chem. 

Rev., 87, 381, 1987. 

177. Liu, S. and Edwards, D. S., Stabilization of (90)y-labeled DOTA-biomolecule conjugates using 

gentisic acid and ascorbic acid, Bioconjug. Chem., 12, 554, 2001. 

178. Mitra, A. et al., Targeting tumor angiogenic vasculature using polymer-RGD conjugates, J. Control. 

Release, 102, 191, 2005. 

179. Koivunen, E., Wang, B., and Ruoslahti, E., Phage libraries displaying cyclic peptides with different 

ring sizes: Ligand specificities of the RGD-directed integrins, Biotechnology (NY), 13, 265, 1995. 

180. Assa-Munt, N. et al., Solution structures and integrin binding activities of an RGD peptide with two 

isomers, Biochemistry, 40, 2373, 2001. 

181. Mitra, A., et al., Polymeric conjugates of mono- and bi-cyclic alphaVbeta3 binding peptides for 

tumor targeting, J. Control. Release, 114, 175, 2006. 

182. Kok, R. J. et al., Preparation and functional evaluation of RGD-modified proteins as alpha(v)beta(3) 

integrin directed therapeutics, Bioconjug. Chem., 13, 128, 2002. 

183. Schraa, A. J. et al., Targeting of RGD-modified proteins to tumor vasculature: A pharmacokinetic 

and cellular distribution study, Int. J. Cancer, 102, 469, 2002. 

184. Xiong, X. B. et al., Intracellular delivery of doxorubicin with RGD-modified sterically stabilized 

liposomes for an improved antitumor efficacy: In vitro and in vivo, J. Pharm. Sci., 94, 1782, 2005. 

185. Dubey, P. K. et al., Liposomes modified with cyclic RGD peptide for tumor targeting, J. Drug 

Target, 12, 257, 2004. 

186. Bibby, D. C. et al., Pharmacokinetics and biodistribution of RGD-targeted doxorubicin-loaded 

nanoparticles in tumor-bearing mice, Int. J. Pharm., 293, 281, 2005. 

187. Seymour, L. W. et al., Effect of molecular weight (Mw) of N-(2-hydroxypropyl)methacrylamide 

copolymers on body distribution and rate of excretion after subcutaneous, intraperitoneal, and intravenous 

administration to rats, J. Biomed. Mater. Res., 21, 1341, 1987. 


184 


188. Seymour, L. W. et al., Influence of molecular weight on passive tumour accumulation of a soluble 

macromolecular drug carrier, Eur. J. Cancer, 31A, 766, 1995. 

189. Kissel, M. et al., Synthetic macromolecular drug carriers: Biodistribution of poly[(N-2-hydroxypropyl)methacrylamide] 

copolymers and their accumulation in solid rat tumors, PDA J. Pharm. Sci. 

Technol., 55, 191, 2001. 

190. Lammers, T. et al., Effect of physicochemical modification on the biodistribution and tumor accumulation 

of HPMA copolymers, J. Control. Release, 110, 103, 2005. 

191. Mitra, A. et al., Technetium-99m-Labeled N-(2-hydroxypropyl) methacrylamide copolymers: 

Synthesis, characterization, and in vivo biodistribution, Pharm. Res., 21, 1153, 2004. 

192. Ulbrich, K. et al., HPMA copolymers with pH-controlled release of doxorubicin: In vitro cytotoxicity 

and in vivo antitumor activity, J. Control. Release, 87, 33, 2003. 

193. Tucker, G. C., Alpha v integrin inhibitors and cancer therapy, Curr. Opin. Investig. Drugs, 4, 722, 

2003. 

194. Porkka, K. et al., A fragment of the HMGN2 protein homes to the nuclei of tumor cells and tumor 

endothelial cells in vivo, Proc. Natl. Acad. Sci. U.S.A., 99, 7444, 2002. 

195. Christian, S. et al., Nucleolin expressed at the cell surface is a marker of endothelial cells in 

angiogenic blood vessels, J. Cell Biol., 163, 871, 2003. 

196. Hoffman, J. A. et al., Progressive vascular changes in a transgenic mouse model of squamous cell 

carcinoma, Cancer Cell, 4, 383, 2003. 

197. Arap, W. et al., Targeting the prostate for destruction through a vascular address, Proc. Natl. Acad. 

Sci. U.S.A., 99, 1527, 2002. 

198. Joyce, J. A. et al., Stage-specific vascular markers revealed by phage display in a mouse model of 

pancreatic islet tumorigenesis, Cancer Cell, 4, 393, 2003. 

199. Ogawa, M. et al., Direct electrophilic radiofluorination of a cyclic RGD peptide for in vivo alpha(v)beta3 

integrin related tumor imaging, Nucl. Med. Biol., 30, 1, 2003. 

200. Su, Z. F. et al., In vitro and in vivo evaluation of a Technetium-99m-labeled cyclic RGD peptide as a 

specific marker of alpha(V)beta(3) integrin for tumor imaging, Bioconjug. Chem., 13, 561, 2002. 

201. Chen, X. et al., Integrin alpha v beta 3-targeted imaging of lung cancer, Neoplasia, 7, 271, 2005. 

202. Chen, X. et al., Synthesis and biological evaluation of dimeric RGD peptide–paclitaxel conjugate as 

a model for integrin-targeted drug delivery, J. Med. Chem., 48, 1098, 2005. 

203. Xiong, X. B. et al., Enhanced intracellular uptake of sterically stabilized liposomal Doxorubicin 

in vitro resulting in improved antitumor activity in vivo, Pharm. Res., 22, 933, 2005. 

204. Thompson, B. et al., Neutral postgrafted colloidal particles for gene delivery, Bioconjug. Chem., 16, 

608, 2005. 

205. Janssen, M. et al., Improved tumor targeting of radiolabeled RGD peptides using rapid dose fractionation, 

Cancer Biother. Radiopharm., 19, 399, 2004. 

206. Bernard, B. et al., Radiolabeled RGD-DTPA-Tyr3-octreotate for receptor-targeted radionuclide 

therapy, Cancer Biother. Radiopharm., 19, 173, 2004. 

207. Capello, A. et al., Increased cell death after therapy with an Arg-Gly-Asp-linked somatostatin analog, 

J. Nucl. Med., 45, 1716, 2004. 


10 

CONTENTS 

Poly(L-Glutamic Acid): 

Efficient Carrier of Cancer 

Therapeutics and Diagnostics 

Guodong Zhang, Edward F. Jackson, Sidney Wallace, and 

Chun Li 

10.1 Introduction ....................................................................................................................... 185 

10.2 Synthesis and Properties of PG......................................................................................... 186 

10.2.1 Chemistry ............................................................................................................ 186 

10.2.2 Biodegradation and Biodistribution .................................................................... 187 

10.3 PG-Anti-Cancer Drug Conjugates .................................................................................... 187 

10.3.1 Anthracyclines ..................................................................................................... 187 

10.3.2 Antimetabolites ................................................................................................... 188 

10.3.3 DNA-Binding Drugs ........................................................................................... 188 

10.4 PG–TXL and PG–CPT: From the Laboratory to the Clinic ............................................ 189 

10.4.1 PG–TXL .............................................................................................................. 189 

10.4.2 PG–CPT............................................................................................................... 192 

10.5 Combination of PG–TXL with Radiotherapy................................................................... 193 

10.6 PG as a Carrier of Diagnostic Agents .............................................................................. 193 

10.6.1 Magnetic Resonance Imaging ............................................................................. 193 

10.6.2 Near-Infrared Fluorescence Optical Imaging ..................................................... 195 

10.7 Conclusions ....................................................................................................................... 195 


References .................................................................................................................................... 196 


Most anti-cancer chemotherapeutic drugs used clinically are limited by a relatively low therapeutic 

index, owing to toxic side effects. 1–3 Over the past several decades, two strategies of improving the 

therapeutic efficacy of anti-cancer agents have emerged. The first approach is the design and development 

of agents that can modulate the molecular processes and pathways specifically associated with 

tumor progression. The success of this approach is shown by the successful introduction of a new 

breed of molecularly targeted anti-cancer agents such as imatinib mesylate (Gleevec), gefitinib 

(Iressa), trastuzumab (Herceptin), and cetuximab (C225, Erbitux). Alternatively, existing 

anti-cancer agents can be made more effective by using delivery systems that bring more drug 

molecules to the tumor site when compared with conventional formulation while reducing exposure 


185

186 

of normal tissues to the drug. In this context, there have been important milestones with the use of 

polymer–drug conjugates in their own right. In particular, poly(L-glutamic acid) (PG)–paclitaxel 

(TXL) (CT2103, XYOTAX) has advanced to phase III clinical trials, and PG–camptothecin (CPT) 

(CT2106) has been tested in phase II clinical trials. The chemistry and applications of PG and its 

conjugates with various chemotherapeutic agents were previously reviewed. 1 In this chapter, the 

applications of PG in the delivery of both chemotherapeutic and diagnostic agents will be updated. 

The results of clinical studies of XYOTAX and CT2106 will also be summarized. Previous chapters 

in this book discuss, in depth, the rationale for the use of polymeric conjugates as cancer therapeutics. 

10.2 SYNTHESIS AND PROPERTIES OF PG 

PG anti-cancer drug conjugates consist of three parts: the PG backbone, the anti-cancer drug, and 

spacers that link the drug molecules with the PG backbone. Both the properties of PG (i.e., the 

molecular weight and its distribution) and the selection of spacers can directly affect the pharmacokinetics 

at both the whole-body and cellular level. For this reason, the synthetic chemistry of the 

preparation of PG will first be discussed. A summary of the physicochemical properties of PG 

conjugates will follow. 

10.2.1 CHEMISTRY 


PG is usually prepared by removing the benzyl protecting group of poly(g-benzyl-L-glutamate) that 

is attained by polymerization of g-benzyl-L-glutamate N-carboxyanhydride (NCA) monomer. 4 

Direct phosgenation of g-benzyl-L-glutamate produces the corresponding NCA monomer. 5,6 In 

this reaction, g-benzyl-L-glutamate is suspended in a dry inert solvent such as ethyl acetate, 

dioxane, or tetrahydrofuran, and it is allowed to heterogeneously react with the cyclizing reagent 

that is normally triphosgene. Researchers have used both the protic and aprotic initiators in the 

polymerization of g-benzyl-L-glutamate NCA monomer. 7 Protic initiators such as primary amines 

are acylated by attack on the 5-position of the NCA, whereas aprotic initiators such as tertiary 

amines and alkoxides act as general bases. The ring-opening polymerization of NCA initiated by 

amines usually produces polymers with relatively broad molecular weight distributions. To circumvent 

this problem, Deming 8 developed the zero-valence nickel catalyst bipyNi(COD) (bipy: 

2,2 0 -bipyridyl; COD: 1,5-cyclooctadiene) to initiate polymerization of NCAs. The active sites 

derived with this initiator are less accessible to side reactions when compared with that derived 

with amine initiators because of steric and electronic effects, resulting in polymers with a narrow 

molecular weight distribution (Mw/Mn, 1.05–1.15). 

Impurities in NCAs can have a detrimental effect on the reproducibility of polyamino acid 

synthesis. NCAs are usually subjected to repeat recrystallization to eliminate trace amounts of 

impurities. A recent report described an efficient method of removal of hydrogen chloride and the 

hydrochloride salt of unreacted starting amino acids. 9 This method consists of extraction of NCA 

solution in ethyl acetate with water and an aqueous alkali solution at 08C. Using highly purified 

monomer NCAs, Aliferis et al. 10 were able to achieve living polymerization initiated with primary 

amine initiators. This technique produced poly(g-benzyl-L-glutamate) with an average molecular 

weight as high as 1.66!10 5 Da and with a relatively narrow molecular weight distribution (Mw/ 

Mn, 1.40). 

The benzyl protecting group in poly(g-benzyl-L-glutamate) is removed by treatment with 

hydrogen bromide. Researchers have also used alkaline hydrolysis of poly(L-methyl glutamate) 

to prepare PG; however, racemization occurred during this process. 11 To avoid using harsh conditions 

for the removal of protecting groups, investigators have prepared glutamic acid NCAs 

having ester protecting groups that are labile under mild acidic conditions. For example, PG can 

be readily obtained from poly(g-piperonyl-L-glutamate) by treating the polymer with trifluoroacetic 

acid. 12 


Poly(L-Glutamic Acid): Efficient Carrier of Cancer Therapeutics and Diagnostics 187 

PG-based block copolymers can be prepared using macroinitiators or polymer coupling reactions. 

The block copolymer PG–vinylsulfone–polyethylene glycol (PEG) was prepared by directly reacting a 

heterofunctional PEG, vinylsulfone–PEG–N-hydroxysuccinimide (NHS), with an amino group at one 

end of the PG chain. 13 Also, Nishiyama et al. 14 used the macroinitiator PEG–NH 2 to initiate polymerization 

of g-benzyl-L-glutamate NCA to produce a PEG–PG block copolymer. 

Another polymer derived from poly(g-benzyl-L-glutamate) is poly(hydroxyethylglutamate) 

(PHEG) that is also a suitable candidate drug carrier. 15 PHEG is a water-soluble, neutral 

polymer that can be readily prepared from poly(g-benzyl glutamate) via aminolysis with 

2-aminoethanol using 2-hydroxypyridine as a catalyst. 16 

The linkage of polymer–drug conjugates is important to the in vivo release of a drug that should 

be stable during circulation but should also allow drug release at an appropriate rate at the tumor 

tissue. 17 Generally, two kinds of linkage are involved in anti-cancer drug delivery by using 

polymer–drug conjugates: enzymatic hydrolysis and nonenzymatic hydrolyzable linkage. Many 

proteolytic enzymes such as cysteine proteinase cathepsins are expressed on the surface of metastatic 

tumor cells. 18,19 The linker chemistry can be designed in such a way that the active agents 

would be released from the polymer–drug conjugates only upon exposure to the proteinases in the 

tumors. Alternatively, hydrolytically labile, pH-sensitive linkages (i.e., hydrozone, ester, acetals, 

etc.) have exhibited their utility for tumorotropic or lysosomotropic delivery because the microenvironments 

of solid tumors are known to be acidic. 20,21 PG–TXL conjugate releases TXL through 

both mechanisms: backbone degradation mediated through proteolytic enzymes and side chain 

hydrolysis of the ester bond formed between glutamic acid and TXL. 

10.2.2 BIODEGRADATION AND BIODISTRIBUTION 

The enzymatic degradability of PG–drug conjugates is influenced by their structure, composition, 

and charge as well as the physicochemical properties of the drug attached to PG. 22–25 PG is more 

susceptible to lysosomal degradation than are poly(aspartic acid) and poly(D-glutamic acid). 26 

Cysteine proteases, particularly cathepsin B, play key roles in the lysosomal degradation of 

PG. 22,27 Although researchers have identified oligomeric glutamic acids as the primary degradation 

products of PG, 28 more recent results demonstrated that monomeric L-glutamic acid is produced in 

the lysosomal degradation of PG. 29 Degradation of the PG backbone may not necessarily lead to the 

release of free drug in every case. 

The biodistribution of PG and its drug conjugates depends on the molecular weight of PG. 

Polymers with a molecular weight lower than the renal clearance threshold are rapidly removed 

from the blood circulation by glomerular filtration. McCormick-Thomson and Duncan 27 found that 

the biodistribution of copolymers of glutamic acid and other hydrophobic amino acids was markedly 

influenced by the composition of the copolymers. These results suggest that the biodistribution 

of PG–drug conjugates is a function of the drugs used, the degree of modification, and the molecular 

weight of PG. 

10.3 PG-ANTI-CANCER DRUG CONJUGATES 

Since doxorubicin (Adriamycin [ADR]) was conjugated with PG in the 1980s, 30 investigators have 

studied various anti-cancer drugs such as anthracyclines, antimetabolites, DNA-binding drugs, 

TXL, and CPT. In particular, researchers have evaluated PG–TXL in phase III clinical trials and 

PG–CPTin phase II clinical trials. 

10.3.1 ANTHRACYCLINES 

The anthracycline group has been one of the most extensively studied groups of drugs used in 

polymer–drug conjugate delivery systems, especially in the 1980s. Based on the assumption that 


188 

greater selectivity toward tumors over normal tissues can be achieved if PG–ADR conjugates only 

degrade and release ADR after being endocytosed by tumor cells, Van Heeswijk and colleagues 30 

investigated three different PG conjugates containing ADR where the amine group in ADR was 

bound to the side chain carboxyl groups of high-molecular-weight PG either directly (with an amide 

bond) or indirectly through GlyGly and GlyGlyGlyLeu spacers, respectively. They investigated the 

degradability of the conjugates mediated by lysosomal enzymes and the subsequent release of ADR 

or ADR-peptide products with low molecular weights using reverse-phase high-performance liquid 

chromatography. The total amount of ADR released after 77 h of incubation was 3.6% for PG-Gly- 

GlyGlyLeu-ADR; 1.0% for PG-GlyGly-ADR; and 0.5% for PG–ADR, suggesting the importance 

of introducing an enzymatically degradable spacer between PG and the drug. In vitro, these 

conjugates exhibited reduced cytotoxicity against L1210 leukemia cells when compared with 

free ADR. In vivo, animals treated with the polymer conjugate containing the tetrapeptide 

spacer showed a similar mean survival duration as compared to those treated with free ADR, 

whereas the conjugate without a peptide spacer was completely inactive. 28,31 

Some have used hydrolytically labile ester bonds 32 and hydrozone bonds 33 to couple doxorubicin 

or daunorubicin (Dau) with PG. In these studies, the ester bond was formed by a reaction of 

14-bromo-daunorubicin with the carboxylic group of PG via a nucleophilic substitution reaction in 

an alkaline aqueous medium. With intravenous administration of these conjugates into mice 

bearing MS-2 sarcoma or Gross’ leukemia, these investigators found that the drug conjugates’ 

potency and efficacy were correlated with the molecular weight of the carrier. For example, 

anti-tumor activity improved when the molecular weight increased from 14,000 to 60,000 Da at 

an equivalent doxorubicin dose of 30 mg/kg. These results can be attributed to the increased 

circulation times of conjugates with high molecular weights. Condensing the methylketone in 

Dau with hydrazide-derived PG yielded PG-hydrazone–Dau conjugates. The acid-sensitive conjugates 

were less cytotoxic than free Dau was in vitro against mouse lymphoma cells. However, these 

conjugates showed significant anti-tumor activity when intravenously injected into mice bearing 

intraperitoneally inoculated Yac lymphoma. 33 

10.3.2 ANTIMETABOLITES 

To enhance the efficacy of 1-b-D-arabinofuranosylcytosine (ara-C) in simple dosage schedules, 

researchers synthesized two ara-C conjugates with PG via an amide bond: one where ara-C is 

directly coupled with the N-4 of ara-C to the carboxyl groups of PG and one where ara-C is linked 

with PG via the aminoalkylphosphoryl side chain introduced at the C-5 0 of ara-C. 34 Both conjugates 

exhibited markedly decreased cytotoxicity in L1210 murine leukemia cells when compared with 

free ara-C. However, the anti-tumor activity of both conjugates was greater than or equal to that of 

free ara-C in mice bearing L1210 tumors after a single intraperitoneal injection of the conjugate. 

The authors suggested that both slow cleavage of free ara-C from the conjugates and protection of 

ara-C from deactivation by cytidine deaminase contributed to the enhanced anti-tumor activity of 

PG–ara-C conjugates in vivo. 

10.3.3 DNA-BINDING DRUGS 


Investigators have conjugated several DNA-binding drugs, including cyclophosphamide, 35 

L-phenylalanine mustard (melphalan), 36 mitomycin C (MMC), 37–39 and cis-dichlorodiammineplatinum 

(II) (cisplatin [CDDP]) 40,41 with PG. Moromoto et al. 36 prepared PG–melphalan by coupling 

the amine group of melphalan with the side-chain carboxyl group of PG in the presence of a watersoluble 

carbodiimide. 36 They investigated the anti-tumor activity of this conjugate in rats bearing 

sarcoma induced by subcutaneous inoculation of sarcoma cells. The authors used PG– 3 H-phenylalanine 

as a model compound to demonstrate that free drug could be released from the conjugate 

and that the conjugate had a tendency to be absorbed through lymphatic routes when compared with 



free 3 H-phenylalanine after subcutaneous administration. 36 Roos et al. 37 conjugated MMC with PG 

by using the aziridine amine of MMC. 37,38 They found that the release rate, in vitro cytotoxicity, 

and in vivo anti-tumor activity of PG–MMC conjugates were influenced by the extent of MMC 

substitution. Furthermore, Seymour et al. 42 conjugated MMC with PHEG using an oligopeptide 

spacer and investigated the effect of this oligopeptide structure on the rate of drug release to 

optimize the MMC–oligopeptide–PHEG conjugates. The rate of MMC release from the conjugates 

was affected by the structure of the oligopeptide spacer; spacers bearing a terminal Gly had the 

fastest release of MMC when compared with other peptide spacers. Another study showed a 

correlation between the in vitro cytotoxicity of PG–MMC conjugates against B16F10 melanoma 

and C26 colorectal carcinoma cells and their hydrolytic stability. 38 Significant in vivo activity of 

the conjugates in mice bearing P338 leukemia or C26 colorectal carcinoma seems to result from 

both the stability of the conjugates in the blood and the rapid drug release, owing to combined 

chemical and enzymatic hydrolysis. 

The clinical use of CDDP is limited by significantly toxic side effects such as acute nephrotoxicity 

and chronic neurotoxicity and a low therapeutic index. To reduce its toxicity and to 

maintain prolonged drug activity, investigators complexed CDDP with PG. 40,41 They found that 

PG–CDDP complex (60 mol CDDP/mol PG; molecular weight, 40,000 Da) was more thermodynamically 

stable and has reduced systemic toxicity when compared with CDDP. PG–CDDP 

was effective in suppressing the growth of OVCAR-3 human ovarian carcinoma in athymic 

mice and showed a broader therapeutic dose range (80% survival at 3–12 mg/kg) when compared 

with the narrow, inconsistent effective dose range of free CDDP (80% survival at 1.0–2.5 mg/kg), 

suggesting that the therapeutic index of CDDP can be improved by complexing it with PG. 

10.4 PG–TXL AND PG–CPT: FROM THE LABORATORY TO THE CLINIC 

10.4.1 PG–TXL 

Taxanes are a class of the widely used and clinically active cytotoxic agents that exert their action 

by promoting tubulin polymerization and microtubule assembly. 43 Because of its poor aqueous 

solubility, use of TXL requires Cremophor and ethanol mixture as a vehicle, and it must be infused 

over 3–24 h. The conjugate PG–TXL where TXL is attached to PG via ester bonds has demonstrated 

significantly enhanced anti-tumor efficacy and improved safety when compared with TXL in 

preclinical studies. 44 

The release of TXL from PG–TXL and the metabolism of PG–TXL in both in vitro and in vivo 

models have been investigated. 45 These studies showed that when PG–TXL was incubated in 

buffered saline or plasma (mouse or human) for 24 h at 378C, less than 14% of the bound TXL 

was released, suggesting that PG–TXL is relatively resistant to plasma esterase. 

Tumor uptake of TXL and PG–TXL was compared using [ 3 H]TXL and PG–[ 3 H]TXL. 46 When 

free [ 3 H]TXL was incubated with MDA-MB453 cells, [ 3 H]TXL was taken up rapidly by the cells 

and was followed by a rapid drug efflux process. In contrast, when the tumor cells were incubated 

with PG–[ 3 H]TXL, a slower uptake and more persistent retention of radioactivity was observed. 

These data suggest that a fraction of PG–[ 3 H]TXL was taken up by tumor cells, possibly through 

pinocytosis, and that [ 3 H]TXL was more readily pumped out of the cells than were the more 

hydrophilic PG–[ 3 H]TXL and its degradation products. A recent report confirmed monoglutamyl-2 

0 -TXL and diglutamyl-2 0 -TXL as the major intracellular metabolites of PG–TXL. 47 

Hydrolysis of these metabolites led to release of free TXL. Specific enzyme inhibitors such as 

CA-074 methyl ester, a cell-permeable irreversible inhibitor of cathepsin B, and EST, a cellpermeable 

irreversible inhibitor of cysteine protease, decreased the formation of monoglutamate 

TXL and free TXL in a tumor cell line that had been incubated with PG–TXL. All of these findings 

are consistent with the results of proteolysis of the PG backbone through the action of cellular 

dipeptidases. Another experiment showed that the metabolism of PG–TXL in non tumor-bearing 


190 


cathepsin B homozygous knockout mice is reduced, but not eliminated; this seems to confirm that 

in addition to cathepsin B, other cysteine proteases on the surface of tumor cells play important 

roles in the release of TXL from PG–TXL. 47 

PG–TXL’s anti-tumor activity was first assessed in a variety of syngeneic and xenogeneic 

tumor models. For example, the maximum tolerated dose of PG–TXL after a single intravenous 

injection in rats and mice was 60 and 160 mg/kg, respectively. 44 In comparison, the maximum 

tolerated dose of TXL in rats and mice was 20 and 60 mg/kg, respectively. Therefore, PG–TXL’s 

use represented a twofold and threefold improvement in toxicity in rats and mice, respectively. To 

determine if PG–TXL has a broad spectrum of anti-tumor activity, its therapeutic activity was 

evaluated against four syngeneic murine tumor cell lines (MCa-4 breast carcinoma, MCa-35 breast 

carcinoma, HCa-1 hepatocarcinoma, and FSa-II sarcoma) intramuscularly inoculated into C3Hf/ 

Kam mice. 48 The anti-tumor and antimetastatic activities of PG–TXL were investigated using 

intraperitoneal injection of the SKOV3ip1 human ovarian tumor cell line in nude mice 49 and 

human MDA-MB-435-Lung2 breast tumors grown in the mammary fat pads of nude mice. 48 

Treatment with PG–TXL exhibited significantly better anti-tumor activity than did treatment 

with TXL alone in all of the tumor models. 

The pharmacokinetic profile and tissue distribution of PG–TXL were examined to verify the 

enhanced permeability and retention effect of macromolecules. 45 It was found that PG–[ 3 H]TXL 

has a much longer half-life in plasma (317 min) than [ 3 H]TXL does (29 min). Consequently, the 

area under the tissue-concentration-time curve (AUC) in tumors was five times greater when mice 

were injected with PG–TXL than when they were injected with TXL. Therefore, enhanced tumor 

uptake and sustained release of TXL from PG–TXL in tumor tissue seem to be among one of the 

major factors contributing to PG–TXL’s markedly improved anti-tumor activity in vivo. 

In the XYOTAX formulation used in clinical studies, the median molecular weight of PG–TXL 

is 48,000 Da, and the content of TXL is about 37% by weight, equivalent to approximately one 

TXL molecule for every 11 glutamic acid units in each PG polymer chain. 47 This formulation 

eliminated the use of Cremophor and alcohol, and it allows infusion of TXL over 30 min. In initial 

clinical trials in the United Kingdom, the investigators gave PG–TXL (CT-2103) to cancer patients 

in a 30-min infusion every three weeks at doses ranging from 30 to 720 mg/m 2 . 50 They detected 

CT-2103 in all the patients’ plasma and observed a long plasma half-life of up to 185 h. Importantly, 

the peak plasma concentration of released free TXL was less than 0.1 mM 24hafter 

administration of CT-2103 at doses up to 480 mg/m 2 (176 mg/m 2 TXL equivalent). As shown in 

Figure 10.1, the plasma concentrations of PG–TXL biphasically declined. The distribution phase 

was prolonged, and the apparent monoexponential terminal phase associated with drug elimination 

appeared approximately 48 h after administration of PG–TXL. The plasma concentration of 

PG–TXL in the terminal phase declined slowly, and the drug could be detected in plasma three 

weeks after administration at a dose of 200 mg/m 2 . In comparison, the plasma concentration of free 

TXL paralleled with the PG–TXL concentration. Importantly, this study found that the AUC of 

free TXL was about 1–2% of the AUC of PG–TXL that supported the in vivo stability of PG–TXL 

in the plasma and the slow, prolonged release of the active moiety. 51 The steady-state volume of 

distribution ranged from 1.3 to 5.9 l/m 2 and was low that suggested that the distribution of PG–TXL 

was restricted mainly to the plasma and other extracellular body fluids. 47 

More than 400 patients have received PG–TXL in phase I and II trials. 52,53 Drug-related events 

that were reported in 10–20% of the patients included thrombocytopenia, diarrhea, leukopenia, 

myalgia, arthralgia, and anemia. Compared with conventional TXL-based treatment, PG–TXL 

showed three safety-related advantages. First, alopecia was rare, and complete hair loss was not 

observed. Second, nausea and vomiting were uncommon. Third, hypersensitivity reactions were 

rarely observed, and those that did occur were usually mild to moderate; therefore, routinely used 

prophylactic premedications were not required. The incidence of significant hypersensitivity 

reactions was less than 1% with no grade 4 hypersensitivity reactions. 



Plasma concentration (µg/ml) 

10 6 

10 5 

10 4 

10 3 

10 2 

10 

1 

(n = 15) 

(n = 15) 

(n = 15) 

(n = 15) 

(n = 13) 

Free paclitaxel 

0 5 10 

Average dose 496 mg 

Conjugates taxanes 

(n = 14) 

(n = 12) 

Phase II trials CT-2103, including the multicenter open-label studies (CTI-1071 and CTI- 

1069), have been completed. 47,53 The former study aimed to determine the rate of response and 

time to disease progression in a heterogeneous population of patients with advanced epithelial 

ovarian cancer who received CT-2103. 47,53 Ninety-nine patients registered in this trial received 

PG–TXL at a dose of 175 mg/m 2 every three weeks. The toxic effects were mild in this heavily 

pretreated population and consisted of the following: grade 3 neuropathy (nZ15), grade 3 neutropenia 

(nZ10), and grade 4 neutropenia (nZ4). Of 18 patients who received one or two prior 

chemotherapy regimens and had CDDP-sensitive disease, five (28%) had a response, and six (33%) 

had stable disease (SD). The cancer was difficult to treat in 21 patients with platinum-resistant 

disease who underwent pretreatment, yet the investigators observed responses in two patients 

(10%) and SD in four patients (19%). In CTI-1069, 47 the researchers aimed to evaluate the efficacy 

and tolerability of PG–TXL in patients with nonsmall cell lung cancer who were 70 years of age or 

older and had an Eastern Cooperative Oncology Group (ECOG) performance status of 2 (PS2). 

Thirty patients were registered in this trial, and they received PG–TXL at a dose of 175 mg/m 2 

(nZ28) or 235 mg/m 2 (nZ2) every three weeks. Two patients had a partial response, whereas 16 

patients had SD lasting at least 10 weeks. Among the 28 patients who received PG–TXL at 175 mg/ 

m 2 , the median survival duration was 8.1 months in those with an ECOG performance status of 0 

(PS0) or 1 (PS1) and 5.4 months in PS2 patients. Both of the patients who received PG–TXL at 

235 mg/m 2 died within 30 days after treatment; one died of neutropenia, whereas the other died of 

septic shock and renal failure. 

More than 700 patients with lung cancer have participated in phase III trials 

PG–TXL(XYOTAX) that include two phase III trials of XYOTAX as first-line treatment in PS2 

patients (STELLAR 3 and STELLAR 4) and one phase III trial of XYOTAX as second-line 

treatment in PS0, PS1, and PS2 patients (STELLAR 2). 47 The total number of patients in the 

STELLAR 2, STELLAR 3, and STELLAR 4 trials was 850, 400, and 477, respectively. Cell 

Therapeutics Inc. (CTI) announced the results of these trials in March and May 2005. Although 

none of the three trials met their primary end points, they did demonstrate similar efficacy, reduced 

Day 

15 

(n = 8) 

20 25 

FIGURE 10.1 Plasma pharmacokinetics of PG–TXL and free TXL in cancer patients. The data are from four 

phase I dose-escalation studies using 1-, 2-, and 3-week schedules. In all of the studies, PG–TXL was 

administrated in a short intravenous infusion. (From Singer, J. W., Shaffer, S., and Baker, B., Anti-cancer 

Drugs, 16, 243–254, 2005. With permission.) 


192 

side effects, and more convenient administration of XYOTAX when compared with the control 

drugs (TXL, docetaxel, gemcitabine, and vinorelbine). 

In addition, clinical data from a pooled analysis of CTI’s STELLAR 3 and 4 trials showed that 

in the 198 women treated on those trials, superior survival was observed in those who received 

XYOTAX (pZ0.03). The most notable impact was among women less than 55 years old and 

presumably pre-menopausal who were treated with XYOTAX compared to standard chemotherapy 

(median survival 10.0 vs. 5.3 months, hazard ratioZ0.51, log rank pZ0.038) while a survival 

trend (pZ0.134) was observed in women 55 years of age and older (post-menopausal) (http:// 

www.ctiseattle.com). The favorable anti-tumor activity among women less than 55 years of age 

stimulated the initiation of the PIONEER 1 clinical trial where about 600 PS2 chemotherapy-naïve 

women with advanced stage NSCLC will be recruited. Each study arm of approximately 300 

patients will be randomized to receive either XYOTAX at 175 mg/m 2 or TXL at 175 mg/m 2 

once every three weeks. The primary endpoint is superior overall survival with several secondary 

endpoints including disease control, response rate in patients with measurable disease, time to 

disease progression, and disease-related symptoms. 

CTI also reported the preliminary results of a phase II study of XYOTAX in combination with 

carboplatin for first-line induction and single-agent maintenance therapy for advanced-stage III/IV 

ovarian cancer in May 2005. Among the 82 patients in this study, 98% had a major tumor response, 

85% had a complete response, and 12% had a partial response during induction of the therapy. At a 

dose of 175 mg/m 2 , XYOTAX and carboplatin (AUCZ6) had grade 3/4 side effects, including 

thrombocytopenia (55%), neuropathy (23%), febrile neutropenia (19%), nausea (15%), anemia 

(11%), and vomiting (7%). The investigators found no grade 4 neuropathy, and only 4% patients 

needed dose delay because of neutropenia. CTI and the Gynecologic Oncology Group are collaborating 

on phase III studies for which they expect to enroll about 1550 patients to receive 

treatment with carboplatin and TXL initially followed by XYOTAX for consolidation in half of 

the patients. 

10.4.2 PG–CPT 


The CPTs are a family of synthetic and semisynthetic analogues of 20(S)-CPT that exhibit a broad 

range of anti-cancer activity by inhibiting topoisomerase-1 activity. Two properties of CPT 

compounds limit their therapeutic efficacy in humans: instability of the lactone form because of 

preferential binding of the carboxylate to serum albumin and a lack of aqueous solubility. PG is an 

effective solubilizing carrier of CPT and serves to protect the E-ring lactone structure in CPT. The 

PG–CPT conjugate was prepared by directly coupling the hydroxy group at the C20(S)-position of 

CPT with the carboxylic acid of PG. 54 When given intravenously in four doses every four days at an 

equivalent CPT dose of 40 mg/kg, PG–CPT delayed the growth of established H322 human lung 

tumors subcutaneously grown in nude mice. In mice that received intratracheal inoculation of H322 

cells, the same treatment prolonged the median survival duration in mice by fourfold when 

compared with that in untreated control mice. These results showed that PG was as efficient 

carrier of CPT. 

Studies have systematically investigated the structural effects of the anti-tumor efficacy of CPT, 

including linkers between PG and CPT, the point of attachment of PG on the CPT molecule, the 

polymer molecular weight, and drug loading. 55–57 First, coupling through the 20(S)-hydroxy group 

of CPT with or without a glycine linker yielded the most active conjugates because this site is 

located in close proximity to the lactone ring; therefore, the E-ring lactone is better protected by the 

linked PG chains. Second, increasing the molecular weight of PG from 33 to 50 kDa improved the 

anti-tumor efficacy of PG–CPT, probably because of an increased plasma half-life and reduced 

renal clearance. Third, based on the CPT-equivalent dosing levels, the investigators compared 

various linkers and found that PG-Gly-CPT and PG–(4-O-butyryl)–CPT seemed had the highest 

anti-tumor activity. Moreover, the linkers also affected the maximum drug payload; for example, 



only 15% (by weight) of CPT loading could be achieved for direct conjugation of PG–CPT by the 

ester linkage because of steric hindrance. On the other hand, up to 50% CPT loading could be 

achieved for the case with a glycine as linker. Fourth, the researchers observed improved anti-tumor 

efficacy of PG-Gly-CPT against HT-29 colon cancer and NCI-H460 lung carcinoma with increased 

CPT loading. However, increasing the loading of CPT to 47% resulted in significantly reduced 

solubility. Therefore, they recommended further investigation of PG-Gly-CPT with loading of CPT 

at 30–35%. 

10.5 COMBINATION OF PG–TXL WITH RADIOTHERAPY 

Combining chemotherapy and radiotherapy has significantly improved response and survival rates 

in patients with many solid tumors. Many chemotherapeutic agents can increase the radiosensitivity 

of tumors, potentiating the tumor response to radiation-caused damage. It was hypothesized that 

combining radiotherapy and chemotherapy using a polymer–drug conjugate may lead to a stronger 

radiosensitizing effect than using the drug alone. Irradiation can, in turn, potentiate the tumor 

response to polymer–drug conjugates by increasing tumor vascular permeability and the uptake 

of these conjugates into solid tumors. To test this hypothesis, PG–TXL was used as a model 

polymer–drug conjugate. 58 Administration of PG–TXL delayed the growth of OCa-1 syngeneic 

murine ovarian tumors in C3Hf/Kam mice. However, when PG–TXL was given in combination 

with tumor irradiation, significantly enhanced anti-tumor activity was observed. Using tumor 

growth delay as an end point, enhancement factors ranging from 1.36 to 4.4 were observed; 

these values depended on the doses of PG–TXL and radiation delivered. It was found that complete 

tumor regression occurred with the use of increased radiation doses (O10 Gy) and PG–TXL doses 

(O80 mg/kg equivalent TXL). 59 Similar results were observed in a mammary MCa-4 carcinoma 

model. 60 In contrast, it was found that combined radiotherapy and TXL treatment yielded an 

enhancement factor of less than 1.0 in MCa-4 tumors, indicating that conjugation of TXL with 

PG is necessary to improve that radiosensitization effect of TXL. When the treatment end point was 

tumor cure, enhancement factors as high as 8.4 and 7.2 were observed after fractionated and singledose 

radiotherapy, respectively. 61,62 

To determine if prior irradiation affects tumor uptake of PG–TXL, [ 3 H]PG–TXL was injected 

into mice with OCa-1 ovarian tumors 24 h after local irradiation at 15 Gy. 59 The uptake of 

[ 3 H]PG–TXL in irradiated tumors was 28–38% higher than that in nonirradiated tumors at different 

times after injection of [ 3 H]PG–TXL, suggesting that irradiation increased the accumulation of 

PG–TXL in the tumors. Therefore, the super-synergistic effect of combined radiotherapy and 

PG–TXL-based chemotherapy is partly ascribed to the enhanced permeability and retention 

effect of macromolecules caused by irradiation. 

10.6 PG AS A CARRIER OF DIAGNOSTIC AGENTS 

10.6.1 MAGNETIC RESONANCE IMAGING 

Noninvasive imaging of intravascular compartments is of critical importance in the clinic. Many 

diseases such as infections, ulcers, cardiovascular diseases, and solid tumors involve gross hemorrhaging, 

abnormal vascular growth, and/or vascular occlusion. 63,64 Magnetic resonance imaging 

(MRI) with blood-pool contrast agents can be used to perform minimally invasive angiography, 

assess angiogenesis, quantify and measure the spacing of blood vessels, and measure blood volume 

and flow. 63,64 MRI blood-pool contrast agents are often paramagnetic gadolinium (Gd) chelates of 

high-molecular-weight polymers that are largely retained within the intravascular space during 

MRI. 65 Recent studies showed that tumor vascular permeability measurements obtained using 

polymeric contrast agents enhanced dynamic MRI that correlated with tumor microvessel 

density counts, suggesting that MRI can be used to characterize tumor angiogenic activity. 66,67 


194 


Furthermore, MRI with polymeric contrast agents may permit more accurate grading of tumor 

invasiveness than those with smaller molecular weight contrast agents, such as Gd–diethylenetriaminepentaacetic 

acid (DTPA). 68,69 

Investigators have synthesized various polymeric contrast agents for MRI, including human 

serum albumin, 70 polylysine, 71,72 dextran, 73,74 dendrimers, 75–78 polyamide, 74,79 and grafted copolymers, 

80 and they evaluated them as blood-pool imaging agents. Although most of these agents 

fulfill the criteria for long blood circulation time and high MRI relaxivity, their safety remains 

unestablished. The clinical applications of many current macromolecular contrast agents are 

limited by slow excretion from the body and the potential toxicity of free Gd 3C ions released by 

the metabolism of the contrast agents. 71–73,81,82 For albumin-based products, the possibility of 

immunogenic responses makes the use of serum proteins less attractive. Dendrimer-based bloodpool 

agents have the advantage of extremely narrow molecular weight distributions; however, these 

agents are not biodegradable. 

Ideally, polymeric contrast agents are degraded and cleared from the body after completion of 

MRI. Based on its good biocompatibility and biodegradability, the metal chelator DTPA was 

conjugated with PG, and the physicochemical and imaging properties of the resulting polymeric 

contrast agents were evaluated. 83 One of these agents, PG–Bz–DTPA–Gd, was synthesized from 

PG and monofunctional p-aminobenzyl-DTPA (penta-tert-butyl ester). It was found that PG–Bz– 

DTPA–Gd was readily degraded upon exposure to an aqueous buffered solution containing cathepsin 

B. The T1 relaxivity of PG–Bz–DTPA–Gd at 1.5 T was four times greater than that of smallmolecular-weight 

Gd–DTPA. Figure 10.2 compares the parametric AUC magnetic resonance 

images for signal intensity integrated over 90 s and 10 min after intravenous injection of gadopentetate 

dimeglumine (Magnevist; Gd–DTPA) and PG–Bz–DTPA–Gd. Magnevist (743 Da) is a 

contrast agent that is clinically used. Whereas Magnevist rapidly diffused into the extravascular 

fluid space over 90 s, PG–Bz–DTPA–Gd was largely retained in the blood vessels for up to 10 min. 

Indeed, contrast enhancement of the vascular compartments was still visible 2 h after injection of 

PG–Bz–DTPA–Gd. 83 

In an effort to increase the rate of polymer degradation, Lu et al. 84 prepared PG–cystamine– 

[Gd(III)-1,4,7,10-tetraazacyclododecane-1,4,7,10-tetraacetic acid (DOTA)] (molecular weight of 

PG, 50,000 Da) where the metal chelator DOTA was conjugated with PG using a disulfide bond. 

They showed that glutathione and other endogenous sulfhydryl-containing biomolecules could 

exchange their free SH group with an S–S bond in PG–cystamine–[Gd(III)–DOTA], resulting in 

FIGURE 10.2 (See color insert following page 522.) Comparison of parametric AUC magnetic resonance 

images obtained at 90 s and 10 min after intravenous injection of Magnevist and PG–Bz–DTPA–Gd. Arrow: 

Blood vessel. 



PG 

PG 

Cl 

rapid release of Gd(III)–DOTA from the polymer and subsequent clearance of Gd-containing 

species from the body. Use of PG–cystamine–[Gd(III)–DOTA] produced significant blood-pool 

contrast enhancement on MRI scans of the heart and blood vessels in nude mice bearing OVCAR-3 

human ovarian carcinoma xenograft when compared with use of small-molecular-weight 

contrast agents. 

10.6.2 NEAR-INFRARED FLUORESCENCE OPTICAL IMAGING 

Near-infrared (NIR) optical imaging has unique advantages for diagnostic imaging of solid tumors. 

Specifically, NIR imaging is a potentially safe, noninvasive method of detecting solid tumors as 

NIR light (650- to 900-nm wavelengths) can penetrate several centimeters into tissue. 58,85 The NIR 

dye indocyanine green (ICG), a model diagnostic agent, has been conjugated with the side 

carboxylic acid of branched PG having a PAMAM core and the terminal folic acid group as a 

targeting moiety (Figure 10.3). 86 The resulting conjugate, PAMAM16–PG–(ICG)–folate, exhibited 

selective binding to KB cells (overexpressing folate receptors) but not to SK-Br3 cells that cannot 

express folate receptors. Additionally, this selective binding was partially blocked by free 

folic acid. 


HN 

O 

NH 

(CH2 ) 2 

m 

O 

N 

H 

O 

n 

NH OC (CH2 ) 2 

COOH 

H 

N 

O 

+ 

N 

Br – 

N 

CO 

HN 

SO 3 – 

COOH 

COOH 

Owing to its favorable physicochemical properties, PG has been shown to be an excellent polymeric 

carrier of both diagnostic and therapeutic agents. PG–TXL has become the first PG-based 

polymeric agent to be tested in clinical trials. The newer generation of PG-based polymers will have 

N 

N 

H 2 N 

HN 

N 

N 

Folic acid 

FIGURE 10.3 Structure of PAMAM16–PG–(ICG)–folate containing folic acid molecules at the termini of the 

polymer and dye molecules (ICG-NH 2) at the side chains of the polymer. (From Tansey, W., Ke, S., Cao, X. Y. 

et al., J. Control. Rel., 94, 39–51, 2004. With permission.) 


196 

to meet a number of challenges, including the development of novel polymers with modulated rates 

of degradation, versatile conjugation chemistry allowing site-specific attachment of targeting 

moieties, and polymerization methods that allow accurate control of polymer molecular weights 

and molecular weight distributions. 


This work was supported in part by the National Cancer Institute (R29 CA74819), the Department 

of Defense (BC960384), and the John S. Dunn Foundation. We thank Donald R. Norwood for 

editing the manuscript. 

REFERENCES 


1. Li, C., Poly(L-glutamic acid)-anticancer drug conjugates, Adv. Drug Deliv. Rev., 54, 695–713, 2002. 

2. Duncan, R., The dawning era of polymer therapeutics, Nat. Rev. Drug Discov., 2, 347–360, 2003. 


1818–1822, 2004. 

4. Idelson, M. E. and Blout, E. R., High molecular weight poly(L-glutamic acid): Preparation and 

physical rotation changes, J. Am. Chem. Soc., 80, 4631–4634, 1958. 

5. Fuller, W. D., Verlander, M. S., and Goodman, M., A procedure for the facile synthesis of amino-acid 

N-carboxyanhydrides, Biopolymers, 15, 1869–1871, 1976. 

6. Katakai, R. and Lisuka, Y., An improved method for the synthesis of N-carboxy amino acid 

anhydrides using trichloromethyl chloroformate, J. Org. Chem., 50, 715–716, 1985. 

7. Block, H., Poly(g-benzyl-L-glutamate) and Other Glutamic Acid Containing Polymers, Gordon and 

Breach, New York, 1983. 

8. Deming, T. J., Facile synthesis of block copolypeptides of defined architecture, Nature, 390, 386–389, 

1997. 

9. Poche, D., Moore, M., and Bowles, J., An unconventional method for purifying the N-carboxyanhydride 

derivatives of g-alkyl-L-glutamates, Synth. Commun., 29, 843–854, 1999. 

10. Aliferis, T., Iatrou, H., and Hadjichristidis, N., Living polypeptides, Biomacromolecules, 5, 

1653–1656, 2004. 

11. Hanby, W. E., Waley, S. G., and Watson, J., Synthetic polypeptides. Part II. Polyglutamic acid, 

J. Chem. Soc., 3239–3249, 1950. 

12. Palacios, P., Bussat, P., and Bichon, D., Novel solid-phase synthesis of thiol-terminated-poly(alphaamino 

acid)–drug conjugate, J. Biochem. Biophys. Methods, 23, 67–72, 1991. 

13. Vega, J., Ke, S., Fan, Z. et al., Targeting doxorubicin to epidermal growth factor receptors by sitespecific 

conjugation of C225 to poly(L-glutamic acid) through a polyethylene glycol spacer, Pharm. 

Res., 20, 826–832, 2003. 

14. Nishiyama, N., Okazaki, S., Cabral, H. et al., Novel cisplatin-incorporated polymeric micelles can 

eradicate solid tumors in mice, Cancer Res., 63, 8977–8983, 2003. 

15. De Marre, A. and Schacht, E., Preparation of 4-nitrophenol carborate esters of poly[N-5-(2-hydroxyethyl)-L-glutamine] 

and coupling with bioactive agents, Markromol. Chem., 193, 3023–3030, 

1992. 

16. Romberg, B., Metselaar, J. M., deVringer, T. et al., Enzymatic degradation of liposome-grafted 

poly(hydroxyethyl L-glutamine), Bioconjug. Chem., 16, 767–774, 2005. 

17. D’Souza, A. J. and Topp, E. M., Release from polymeric prodrugs: Linkages and their degradation, 

J. Pharm. Sci., 93, 1962–1979, 2004. 

18. Thomssen, C., Schmit, M., Goretzki, L. et al., Prognostic value of the cysteine proteases cathepsins B 

and cathepsin L in human breast cancer, Clin. Cancer Res., 1, 741–746, 1995. 

19. Mai, J., Waisman, D. M., and Sloane, B. F., Cell surface complex of cathepsin B/annexin II tetramer in 

malignant progression, Biochim. Biophys. Acta, 1477, 215–230, 2000. 

20. Kornguth, S. E., Kalinke, T., Robins, H. I., Cohen, J. D., and Turski, P., Preferential binding of 

radiolabeled poly-L-lysines to C6 and U87 MG glioblastomas compared with endothelial cells 

in vitro, Cancer Res., 49, 6390–6395, 1989. 



21. Hoste, K., De Winne, K., and Schacht, E., Polymeric prodrugs, Int. J. Pharm., 277, 119–131, 2004. 

22. Chiu, H. C., Kopeckova, P., Deshmane, S. S., and Kopecek, J., Lysosomal degradability of poly(alphaamino 

acids), J. Biomed. Mater. Res., 34, 381–392, 1997. 

23. Tsutsumi, A., Perly, B., Forchioni, A., and Charahaty, C., A magnetic resonance study of the 

segmental motion and local conformations of poly(L-glutamic acid) in aqueous solutions, Macromolecules, 

11, 977–986, 1978. 

24. Dolnik, V., Novotny, M., and Chmelik, J., Electromigration behavior of poly(L-glutamate) conformers 

in concentrated polyacrylamide gels, Biopolymers, 33, 1299–1306, 1993. 

25. Miller, W. G., Degradation of synthetic polypeptides II. degradation of poly-L-glutamic acid by 

proteolytic enzymes in 0.20 M sodium chloride, J. Am. Chem. Soc., 86, 3913–3918, 1964. 

26. Kishore, B. K., Kallay, Z., Lambricht, P., Laurent, G., and Tulkens, P. M., Mechanism of protection 

afforded by polyaspartic acid against gentamicin-induced phospholipidosis I. Polyaspartic acid binds 

gentamicin and displaces it from negatively charged phospholipid layers in vitro, J. Pharmacol. Exp. 

Ther., 255, 867–874, 1990. 

27. McCormick-Thomson, L. A. and Duncan, R., Poly(amino acid) copolymers as a potential soluble drug 

delivery system. I: Pinocytotic uptake and lysosomal degradation measured in vitro, J. Bioact. 

Compat. Polym., 4, 242–251, 1989. 

28. Hoes, C. J. T., Potman, W., Van Heeswijk, W. A. R. et al., Optimization of macromolecular prodrugs 

of the antitumor antibiotic Adriamycin, J. Control. Rel., 2, 205–213, 1985. 

29. Singer, J. W., Nudelman, E., De Vries, P., Shaffer, S., Metabolism of poly-L-glutamic acid (PG) 

paclitaxel (CT-2103): Intratumoral peptide cleavage followed by esterolysis, In 5th International 

Symposium on Polymer Therapeutics: From Laboratory to Clinical Practice, Cardiff, UK, 2002. 

30. Van Heeswijk, W. A. R., Hoes, C. J. T., Stoffer, T. et al., The synthesis and characterization of 

polypeptide–adriamycin conjugates and its complexes with Adriamycin. Part 1, J. Control. Rel., 1, 

301–315, 1985. 

31. Hoes, C. J. T., Grootoonk, J., Duncan, R. et al., Biological properties of adriamycin bound to 

biodegradable polymeric carriers, J. Control. Rel., 23, 37–54, 1993. 

32. Zunino, F., Pratesi, G., and Micheloni, A., Poly(carboxylic acid) polymers as carriers for anthracyclines, 

J. Control. Rel., 10, 65–73, 1989. 

33. Hurwitz, E., Wilchek, M., and Pitha, J., Soluble macromolecules as carriers for daunorubicin, J. Appl. 

Biochem., 2, 25–35, 1980. 

34. Kato, Y., Saito, M., Fukushima, H., Takeda, Y., and Hara, T., Antitumor activity of 1-beta-D-arabinofuranosylcytosine 

conjugated with polyglutamic acid and its derivative, Cancer Res., 44, 25–30, 

1984. 

35. Batz, H. G., Ringsdorf, H., and Ritter, H., Pharmacologically active polymers. 7. Cyclophosphamideand 

steroid hormone-containing polymers as potential anticancer compounds, Makromol. Chem., 175, 

2229–2239, 1974. 

36. Moromoto, Y., Sugibayashi, K., Sugihara, S. et al., Antitumor agent poly(amino acid) conjugates as a 

drug carrier in chemotherapy, J. Pharm. Dyn., 7, 688–698, 1984. 

37. Roos, C. F., Matsumoto, S., Takarura, Y., Hashida, M., and Sezaki, H., Physicochemical and antitumor 

characteristics of some polyamino acid prodrugs of mitomycin C, Int. J. Pharm., 22, 75–87, 

1984. 

38. Soyez, H., Schacht, E., and Vanderkerken, S., The crucial role of spacer groups in macromolecular 

prodrug design, Adv. Drug Deliv. Rev., 21, 81–106, 1996. 

39. de Marre, A., Seymour, L. W., and Schacht, E., Evaluation of the hydrolytic and enzymatic stability of 

macromolecular mitomycin C derivatives, J. Control. Rel., 31, 89–97, 1994. 

40. Avichezer, D., Schechter, B., and Arnon, R., Functional polymers in drug delivery: Carrier-supported 

CDDP (cisplatin) complexes of polycarboxylates-effects on human ovarian carcinoma, React. Funct. 

Polym., 36, 59–69, 1998. 

41. Schlechter, B., Neumann, A., Wilchek, M., and Arnon, R., Soluble polymers as carriers of cisplatinum, 

J. Control. Rel., 10, 75–87, 1989. 

42. Seymour, L. W., Soyez, H., De Marre, A., Shoaibi, M. A., and Schacht, E. H., Polymeric prodrugs of 

mitomycin C designed for tumour tropism and sustained activation, Anti-cancer Drug Des., 11, 

351–365, 1996. 

43. Horwitz, S. B., Taxol (Paclitaxel): Mechanisms of action, Ann. Oncol., 5, S3–S6, 1994. 


198 


44. Li, C., Yu, D. F., Newman, R. A. et al., Complete regression of well-established tumors using a novel 

water-soluble poly(L-glutamic acid)–paclitaxel conjugate, Cancer Res., 58, 2404–2409, 1998. 

45. Li, C., Newman, R. A., Wu, Q. P. et al., Biodistribution of paclitaxel and poly(L-glutamic acid)– 

paclitaxel conjugate in mice with ovarian OCa-1 tumor, Cancer Chemother. Pharmacol., 46, 

416–422, 2000. 

46. Oldham, E. A., Li, C., Ke, S., Wallace, S., and Huang, P., Comparison of action of paclitaxel and poly(Lglutamic 

acid)–paclitaxel conjugate in human breast cancer cells, Int. J. Oncol., 16, 125–132, 2000. 

47. Singer, J. W., Shaffer, S., Baker, B. et al., Paclitaxel poliglumex (XYOTAX; CT-2103): An intracellularly 

targeted taxane, Anticancer Drugs, 16, 243–254, 2005. 

48. Li, C., Price, J. E., Milas, L. et al., Antitumor activity of poly(L–glutamic acid)–paclitaxel on 

syngeneic and xenografted tumors, Clin. Cancer Res., 5, 891–897, 1999. 

49. Auzenne, E., Donato, N. J., Li, C. et al., Superior therapeutic profile of poly-L-glutamic acid–paclitaxel 

copolymer compared with taxol in xenogeneic compartmental models of human ovarian 

carcinoma, Clin. Cancer Res., 8, 573–581, 2002. 

50. Todd, R., Sludden, J., Boddy, A. V. et al., Phase I and pharmacological study of CT-2103, a poly(Lglutamic 

acid)–paclitaxel conjugate, Proc. Am. Soc. Clin. Oncol., 20, 2001 (abstr 439). 

51. Bernareggi, A., Oldham, F., Barone, C., Tumor-targeted taxane: PK evidence for an enhanced permeability 

and retention effect with paclitaxel poliglumex (CT-2103), Presented at the European 

Society for Medical Oncology (ESMO), October 29–November 2, 2004, Vienna, Austria. 

52. Markman, M., Improving the toxicity profile of chemotherapy for advanced ovarian cancer: A 

potential role for CT-2103, J. Exp. Ther. Oncol., 4, 131–136, 2004. 

53. Sabbatini, P., Aghajanian, C., Dizon, D. et al., Phase II study of CT-2103 in patients with recurrent 

epithelial ovarian, fallopian tube, or primary peritoneal carcinoma, J. Clin. Oncol., 22, 4523–4531, 

2004. 

54. Zou, Y., Wu, Q. P., Tansey, W. et al., Effectiveness of water soluble poly(L-glutamic acid)–camptothecin 

conjugate against resistant human lung cancer xenografted in nude mice, Int. J. Oncol., 18, 

331–336, 2001. 

55. Singer, J. W., De Vries, P., Bhatt, R. et al., Conjugation of camptothecins to poly(L-glutamic acid), 

Ann. N Y Acad. Sci., 922, 136–150, 2000. 

56. Singer, J. W., Bhatt, R., Tulinsky, J. et al., Water-soluble poly(L-glutamic acid)-Gly-camptothecin 

conjugates enhance camptothecin stability and efficacy in vivo, J. Control. Rel., 74, 243–247, 2001. 

57. Bhatt, R., de Vries, P., Tulinsky, J. et al., Synthesis and in vivo antitumor activity of poly(L-glutamic 

acid) conjugates of 20S-camptothecin, J. Med. Chem., 46, 190–193, 2003. 

58. Hawrysz, D. J. and Sevick-Muraca, E. M., Developments toward diagnostic breast cancer imaging 

using near-infrared optical measurements and fluorescent contrast agents, Neoplasia, 2, 388–417, 

2000. 

59. Li, C., Ke, S., Wu, Q. P. et al., Tumor irradiation enhances the tumor-specific distribution of 

poly(L-glutamic acid)–conjugated paclitaxel and its antitumor efficacy, Clin. Cancer Res., 6, 

2829–2834, 2000. 

60. Ke, S., Milas, L., Charnsangavej, C., Wallace, S., and Li, C., Potentiation of radioresponse by 

polymer–drug conjugates, J. Control. Rel., 74, 237–242, 2001. 

61. Milas, L., Mason, K. A., Hunter, N., Li, C., and Wallace, S., Poly(L-glutamic acid)–paclitaxel conjugate 

is a potent enhancer of tumor radiocurability, Int. J. Radiat. Oncol. Biol. Phys., 55, 707–712, 

2003. 

62. Tishler, R. B., Polymer–conjugated paclitaxel as a radiosensitizing agent-a big step forward for 

combined modality therapy?, Int. J. Radiat. Oncol. Biol. Phys., 55, 563–564, 2003. 

63. Bogdanov, A. A., Lewin, M., and Weissleder, R., Approaches and agents for imaging the vascular 

system, Adv. Drug Deliv. Rev., 37, 279–293, 1993. 

64. Kroft, L. J. and de Roos, A., Blood pool contrast agents for cardiovascular MR imaging, J. Magn. 

Reson. Imaging, 10, 395–403, 1999. 

65. Brasch, R. C., New directions in the development of MR imaging contrast media, Radiology, 183, 

1–11, 1992. 

66. Su, M. Y., Muhler, A., Lao, X., and Nalcioglu, O., Tumor characterization with dynamic contrastenhanced 

MRI using MR contrast agents of various molecular weights, Magn. Reson. Med., 39, 

259–269, 1998. 



67. van Dijke, C. F., Brasch, R. C., Roberts, T. P. et al., Mammary carcinoma model: Correlation of 

macromolecular contrast-enhanced MR imaging characterizations of tumor microvasculature and 

histologic capillary density, Radiology, 198, 813–818, 1996. 

68. Gossmann, A., Okuhata, Y., Shames, D. M. et al., Prostate cancer tumor grade differentiation with 

dynamic contrast-enhanced MR imaging in the rat: Comparison of macromolecular and small-molecular 

contrast media-preliminary experience, Radiology, 213, 265–272, 1999. 

69. Daldrup, H., Shames, D. M., Wendland, M. et al., Correlation of dynamic contrast-enhanced magnetic 

resonance imaging with histologic tumor grade: Comparison of macromolecular and small-molecular 

contrast media, Pediatr. Radiol., 28, 67–78, 1998. 

70. van Bemmel, C. M., Wink, O., Verdonck, B., Viergever, M. A., and Niessen, W. J., Blood pool 

contrast-enhanced MRA: Improved arterial visualization in the steady state, IEEE Trans. Med. 

Imaging, 22, 645–652, 2003. 

71. Schuhmann-Giampieri, G., Schmitt-Willich, H., Frenzel, T., Press, W. R., and Weinmann, H. J., 

In vivo and in vitro evaluation of Gd–DTPA–polylysine as a macromolecular contrast agent for 

magnetic resonance imaging, Invest. Radiol., 26, 969–974, 1991. 

72. Bogdanov, J., Alexei, A., Weissleder, R., and Brady, T. J., Long-circulating blood pool imaging 

agents, Adv. Drug Deliv. Rev., 16, 335–348, 1995. 

73. Wang, S. C., Wikstrom, M. G., White, D. L. et al., Evaluation of Gd–DTPA–labeled dextran as an 

intravascular MR contrast agent: Imaging characteristics in normal rat tissues, Radiology, 175, 

483–488, 1990. 

74. Rebizak, R., Schaefer, M., and Dellacherie, E., Polymeric conjugates of Gd(3C)-diethylenetriaminepentaacetic 

acid and dextran. 1. Synthesis, characterization, and paramagnetic properties, 

Bioconjug. Chem., 8, 605–610, 1997. 

75. Wiener, E. C., Brechbiel, M. W., Brothers, H. et al., Dendrimer-based metal chelates: A new class of 

magnetic resonance imaging contrast agents, Magn. Reson. Med., 31, 1–8, 1994. 

76. Adam, G., Neuerburg, J., Spuntrup, E. et al., Gd–DTPA–cascade-polymer: Potential blood pool 

contrast agent for MR imaging, J. Magn. Reson. Imaging, 4, 462–466, 1994. 

77. Kobayashi, H. M. and Brechbiel, W., Dendrimer-based macromolecular MRI contrast agents: Characteristics 

and application, Mol. Imaging, 2, 1–10, 2003. 

78. Kim, Y. H., Choi, B. I., Cho, W. H. et al., Dynamic contrast-enhanced MR imaging of VX2 carcinomas 

after X-irradiation in rabbits: Comparison of gadopentetate dimeglumine and a 

macromolecular contrast agent, Invest. Radiol., 38, 539–549, 2003. 

79. Unger, E. C., Shen, D., Wu, G. et al., Gadolinium-containing copolymeric chelates—a new potential 

MR contrast agent, Magma, 8, 154–162, 1999. 

80. Gupta, H., Wilkinson, R. A., Bogdanov, A. A., Callahan, R. J., and Weissleder, R., Inflammation: 

Imaging with methoxy poly(ethylene glycol)–poly-L-lysine–DTPA, a long-circulating graft copolymer, 

Radiology, 197, 665–669, 1995. 

81. Rebizak, R., Schaefer, M., and Dellacherie, E., Polymeric conjugates of Gd(3C)-diethylenetriaminepentaacetic 

acid and dextran. 2. Influence of spacer arm length and conjugate molecular mass on the 

paramagnetic properties and some biological parameters, Bioconjug. Chem., 9, 94–99, 1998. 

82. Franano, F. N., Edwards, W. B., Welch, M. J. et al., Biodistribution and metabolism of targeted and 

nontargeted protein-chelate–gadolinium complexes: Evidence for gadolinium dissociation in vitro 

and in vivo, Magn. Reson. Imaging, 13, 201–214, 1995. 

83. Wen, X., Jackson, E. F., Price, R. E. et al., Synthesis and characterization of poly(L-glutamic acid) 

gadolinium chelate: A new biodegradable MRI contrast agent, Bioconjug. Chem., 15, 1408–1415, 2004. 

84. Lu, Z. R., Wang, X., Parker, D. L., Goodrich, K. C., and Buswell, H. R., Poly(L-glutamic acid) 

Gd(III)–DOTA conjugate with a degradable spacer for magnetic resonance imaging, Bioconjug. 

Chem., 14, 715–719, 2003. 

85. Ke, S., Wen, X., Gurfinkel, M. et al., Near-infrared optical imaging of epidermal growth factor 

receptor in breast cancer xenografts, Cancer Res., 63, 7870–7875, 2003. 

86. Tansey, W., Ke, S., Cao, X. Y. et al., Synthesis and characterization of branched poly(L-glutamic acid) 

as a biodegradable drug carrier, J. Control. Rel., 94, 39–51, 2004. 


11 

CONTENTS 

Noninvasive Visualization of 

In Vivo Drug Delivery of 

Paramagnetic Polymer 

Conjugates with MRI 

Zheng-Rong Lu, Yanli Wang, Furong Ye, Anagha 

Vaidya, and Eun-Kee Jeong 

11.1 Introduction ....................................................................................................................... 201 

11.2 Magnetic Resonance Imaging ........................................................................................... 202 

11.2.1 Principles of MRI ................................................................................................ 202 

11.2.2 Contrast-Enhanced MRI...................................................................................... 203 

11.3 MR Imaging of Paramagnetic HPMA Copolymer Conjugates........................................ 204 

11.3.1 MR Imaging of a Paramagnetic HPMA Conjugate in an Animal 

Tumor Model....................................................................................................... 204 

11.3.2 MR Imaging of a Paramagnetic HPMA Conjugate for Bone Delivery............. 206 

11.4 MR Imaging of Paramagnetic Poly(L-Glutamic Acid) Conjugates ................................. 206 

11.4.1 MR Imaging of PGA–(Gd-DOTA) Conjugate in an Animal 

Tumor Model....................................................................................................... 206 

11.4.2 MR Imaging of PGA–Mce6–(Gd-DOTA) Conjugate in an Animal 

Tumor Model....................................................................................................... 208 

11.5 Summary............................................................................................................................ 210 

References .................................................................................................................................... 210 


The conjugation of therapeutics to water-soluble biomedical polymers increases the aqueous solubility 

of hydrophobic drugs, prolongs in vivo drug retention time, reduces systemic toxicity, and 

enhances therapeutic efficacy. 1,2 Polymer drug conjugates can preferentially accumulate in solid 

tumor tissues due to the hyperpermeability of tumor blood vessels or the so-called “enhanced 

permeability and retention (EPR) effect.” 3,4 Polymer drug conjugates can also down-regulate or 

overcome multiple drug resistance. 5,6 Because of these unique properties, polymer drug conjugates 

exhibit higher therapeutic efficacy than the corresponding therapeutics alone. Several polymer drug 

conjugates are currently used in clinical cancer treatment, and more are in the pipeline of clinical 


201

202 

development. For example, styrene–maleic acid copolymer neocarzinostatin conjugate (SMANCS) 

is used for liver cancer treatment in Japan; 7 pegylated adenosine deaminase 8 and asparaginase 9,10 

are used for enzyme replacement therapy for immunodeficiency and for the treatment of acute 

lymphoblastic leukemia, respectively. Many other polymer drug conjugates, including pegylated 

interferon a, 11 PEG–camptothecin conjugate, 12 poly(L-glutamic acid) paclitaxel conjugate, 13 

poly[N-(2-hydroxypropyl)methacrylamide] (PHPMA)-cis-platinate conjugate, 14 and PHPMA– 

doxorubicin conjugate 15,16 are in various phases of clinical trials. 

The in vivo behavior, including pharmacokinetics, biodistribution, drug delivery efficiency and 

therapeutic efficacy, of polymer drug conjugates has been traditionally evaluated by using blood 

and urine sampling, biopsy-based methods, and symptom-based observations. These methods are 

sometimes invasive and cannot accurately provide real time information on the interaction of 

polymers with various organs and tissues, the targeting and delivery efficiency of the drug delivery 

system, and therapeutic response. A large number of animals are also required in preclinical 

development. Sometimes, the data obtained by conventional biopsies may be misleading, which 

might be one of the causes of the efficacy discrepancy between preclinical development in animal 

models and human clinical trials and the failure of some clinical trials due to improper selection of 

drug candidates. 

In vivo drug delivery by polymer drug conjugates involves circulation of the conjugates in the 

blood; interaction with the major organs including the liver, heart, lungs, spleen and kidneys, etc.; 

transport to target tissue; and uptake by target cells and drug release. Direct and continuous 

evaluation of drug delivery efficiency and tissue interaction of polymeric drug conjugates is critical 

to develop more efficacious drug delivery systems. Recent advancements in biomedical imaging 

technology have provided the essential tools for noninvasive and continuous in vivo evaluation for 

drug delivery. Nuclear medicine, including positron emission tomography (PET) and single photon 

emission computed tomography (SPECT), is a clinical imaging modality with high detection 

sensitivity (ca. 10 K9 M). Magnetic resonance imaging (MRI) is a noninvasive clinical imaging 

modality with high spatial resolution. Both imaging modalities are able to noninvasively visualize 

in vivo drug delivery with polymeric drug conjugates. The polymer conjugates can be labeled with 

imaging probes and noninvasively and continuously monitored with the imaging modalities. 

The number of experimental animals can be dramatically reduced in the preclinical development 

of the conjugates. Biomedical imaging will provide real-time information of the in vivo behavior of 

polymeric drug conjugates. In this chapter, we will focus on noninvasive visualization of in vivo 

drug delivery with paramagnetic polymeric conjugates and contrast-enhanced MRI. 

11.2 MAGNETIC RESONANCE IMAGING 

11.2.1 PRINCIPLES OF MRI 


MRI is a clinical imaging modality that produces three-dimensional anatomic images with high 

spatial resolution and provides in vivo physiological properties including physiochemical information, 

flow diffusion, and motion in the tissues. The physics of MRI is complicated and has been 

detailed in a number of monographs. 17,18 The fundamental principles of MRI are briefly reviewed 

in this section. A normal human adult has approximately 60% of the body weight as water. Water 

protons have a magnetic moment and the orientations of the magnetic moments are random in the 

absence of an external magnetic field. When a patient is placed in a MRI scanner, the proton 

magnetic moments align either along or against the static magnetic field (B0) of the scanner and 

create a net magnetization pointing in the direction of the main magnetic field of the scanner. 

The magnitude of the magnetization is proportional to the external magnetic field strength as well 

as the amount of protons. 

Nuclear magnetic resonance (NMR) measures the change of the magnetization by applying 

radiofrequency (RF) pulses. When an RF pulse is applied to create an oscillating electromagnetic 


Noninvasive Visualization of In Vivo Drug Delivery of Paramagnetic Polymer Conjugates 203 

field (B1) perpendicular to the main field, the longitudinal magnetization is tipped away from the 

static magnetic field and deviates from the equilibrium state (in longitudinal space). A perfect 908 

RF pulse flips the total longitudinal magnetization onto the transverse plane. Immediately after 

excitation with the RF pulse, the transverse magnetization then undergoes the decay process. 

The decaying transverse magnetization induces an NMR signal as an electromotive force (emf) 

on an RF coil. Simultaneously, the nuclear spin system approaches toward the equilibrium state 

with a characteristic recovery time T1, which depends upon the physical and chemical environment 

of the molecules. 

Longitudinal T1 relaxation involves the return of protons from the high energy state back to 

equilibrium by dissipating their excess energy to their surroundings. The process is also called spinlattice 

relaxation. Transverse (T 2) relaxation involves the transfer of the spin angular momentum 

among the protons via the interactions such as dipole–dipole interaction. The process is called 

spin–spin relaxation. In biological systems, the T2 relaxation time is typically shorter than T1 

relaxation time. Because the main magnetic field of MRI scanners is not perfectly homogenous, 

the inhomogeneity of the magnetic field also causes dephasing of individual magnetizations of 

protons, resulting in more rapid loss of transverse magnetization. The process is called 

T 2 relaxation. 

In MR imaging, the position dependent frequency and phase are encoded into the transverse 

magnetization, and the measured signals are Fourier transformed to construct the spatial map 

(image). MR images are created based on the proton density, or T 1 and T 2 relaxation rates, and 

flow and diffusion properties in different tissues. These parameters vary between different tissues 

and create image contrast among the tissues. Tissues with a short T1 relaxation time give a strong 

MR signal, resulting in bright images in T1-weighted MRI. Tissues with a short T2 relaxation time 

produce a weak MR signal, resulting in dark images in T2-weighted MRI. 

11.2.2 CONTRAST-ENHANCED MRI 

In many cases, the differences between the MR signal intensities of normal and diseased tissues are 

not large enough to produce obvious contrast for definitive diagnosis. Contrast agents are used to 

enhance the image contrast between normal and diseased tissues to improve diagnostic sensitivity 

and specificity. 19,20 MRI contrast agents are either paramagnetic metal chelates or ultrasmall 

superparamagnetic iron oxides. These agents interact with surrounding water molecules and 

increase the relaxation rates (1/T1 and 1/T2) of water protons. The change of relaxation rate is 

linearly proportional to the concentration of contrast agents, [M]: 

1=T i;obs Z 1=T i;d Cr i½MŠ; (11.1) 

where iZ1,2. In Equation 11.1, 1/T i,obs is the observed relaxation rate with a contrast agent, 1/T i,d 

the relaxation rate without the contrast agent, and ri the relaxivity of the contrast agent. Relaxivity is 

a measurement of the efficiency of a contrast agent to enhance the relaxation rate of water protons 

or the efficiency to generate image contrast enhancement. Because the uptake of contrast agents 

varies between different tissues, MR image contrast is enhanced in the tissue with a high contrast 

agent concentration compared to that with a low concentration. 

MRI contrast agents approved for clinical applications are mainly stable paramagnetic chelates, 

e.g., Gd(III) chelates 21–24 and Mn(II) chelates, 25 and ultrasmall superparamagnetic iron oxide 

(USPIO). 26 Gd(III) chelates of DTPA, DOTA, or their derivatives are commonly used for 

contrast-enhanced MRI. Both Gd(III) and Mn(II) chelates are mainly used as T1 contrast agents, 

whereas USPIO is used as a T2 contrast agent. These agents significantly enhance image contrast 

between normal tissue and diseased tissue and improve the diagnostic accuracy. Currently, contrast 

agents are used in approximately 30% of clinical MRI examinations. Macromolecular Gd(III) 

complexes have also been developed as MRI contrast agents. 27,28 These agents have a prolonged 


204 

blood circulation time and preferentially accumulate in solid tumors due to the EPR effect. Macromolecular 

Gd(III) complexes have demonstrated superior contrast enhancement in MR 

angiography and cancer imaging in animal models, but are not yet approved for clinical 

applications. 

Contrast-enhanced MRI is a useful method for noninvasive in vivo evaluation of polymeric 

drug delivery systems after the systems are labeled with MRI contrast agents. The real-time 

pharmacokinetics and biodistribution of the labeled drug delivery systems can be continuously 

visualized by contrast-enhanced MRI. It has a potential to provide accurate four-dimensional 

information of in vivo properties of the drug delivery systems. 

11.3 MR IMAGING OF PARAMAGNETIC HPMA COPOLYMER CONJUGATES 

N-(2-Hydroxypropyl)methacrylamide (HPMA) copolymer is a biocompatible water-soluble polymeric 

carrier for therapeutics. 1 The conjugation of anti-cancer drugs to HPMA copolymers results 

in many advantageous features over small molecular therapeutics, including improved water solubility 

and bioavailability, preferential accumulation of the conjugates in solid tumors, reduced 

nonspecific toxicity, and down-regulation of multi-drug resistance. 1,29 A number of HPMA copolymer 

drug conjugates have been systematically investigated both in vitro and in vivo. Several 

HPMA copolymer-anti-cancer drug conjugates are now in various phases of clinical development 

for cancer treatment. 30 The incorporation of MRI contrast agents into HPMA copolymer conjugates 

allows direct visualization of real-time drug delivery of the conjugates in animal models with 

contrast-enhanced MRI. The noninvasive imaging reveals some interesting features of drug 

delivery with the conjugates that cannot be obtained with conventional surgery based evaluations. 

11.3.1 MR IMAGING OF A PARAMAGNETIC HPMA CONJUGATE 

IN AN ANIMAL TUMOR MODEL 


HPMA copolymers have been labeled with radioactive probes including 131 Iand 99m Tc for 

noninvasive visualization of in vivo drug delivery of the conjugates with g-scintigraphy. 31–33 

Nuclear medicine is highly sensitive for detecting labeled conjugates, but its poor spatial resolution 

cannot provide detailed biodistribution of conjugates in tissues and organs. In comparison, contrastenhanced 

MRI provides clear visualization of the biodistribution of paramagnetically labeled 

conjugates with high spatial resolution. 

HPMA copolymers were labeled with an MRI contrast agent, Gd-DOTA, and investigated in 

an animal tumor model with contrast-enhanced MRI. 34 The structure of a labeled conjugate, 

poly[HPMA-co-(MA-Gly-Gly-1,6-hexanediamine-(Gd-DOTA))], is shown in Figure 11.1. 

The molecular weight of the polymer conjugate was 25.6 kDa (PDZ1.50) and the Gd content 

was 0.33 mmol Gd/g polymer. The T1 relaxivity of the conjugate was 6.0 mM K1 s K1 per attached 

Gd(III) chelate at 3 T. Figure 11.2 shows the dynamic, contrast-enhanced, three-dimensional 

maximum intensity projection (MIP) (Figure 11.2a) and 2D coronal MR images (Figure 11.2b) 

of a mouse bearing a human prostate carcinoma DU-145 xenograft and a Kaposi’s sarcoma xenograft. 

The conjugate was administrated intravenously at a Gd equivalent dose of 0.1 mmo1/kg body 

weight. The pharmacokinetics and biodistribution of the conjugate were clearly revealed in the 

contrast-enhanced MR images. 

Strong MRI signal was observed in the heart, liver, kidneys and vasculature in the 3D images at 

the initial stage after the injection of the conjugates. The signal intensity decreased over a 60-min 

period; meanwhile, the signal intensity in the urinary bladder increased gradually, indicating that 

the conjugate was excreted via renal filtration. Significant contrast enhancement was still visible in 

the heart at 60 min. After 22 h, most of the conjugate was cleared from the body except the liver, as 

shown in Figure 11.2. The signal intensity in the liver was stronger at 22 h than in the precontrast 

image, suggesting significant interaction of HPMA copolymers with the liver. 



CH2 

CH3 

C 

m 

O 

NH 

CH2 

CH OH 

CH3 

CH 2 

CH3 

C 

n 

O 

NH 

CH2 

C O 

NH 

CH2 

O O 

NH(CH2) 6NH C 

− OOC 

N N 

Gd 

N N 

3+ 

COO − 

COO − 

FIGURE 11.1 The structure of poly[HPMA-co-(MA-Gly-Gly-1,6-hexanediamine-(Gd-DOTA))]. 

The coronal images cross-sectioning the tumor tissues showed heterogeneous uptake of HPMA 

copolymers in two different tumors (Figure 11.2b). The prostate carcinoma had a smaller size, but 

stronger MR signal than the Kaposi’s sarcoma. For Kaposi’s sarcoma, the contrast enhancement 

was mainly observed in the periphery of the tumor tissue and to a lower extent in the inner tumor 

tissue. A plausible explanation is that the interstitium of the large tumor was possibly necrotic, 

which would limit the access of the conjugate. The results suggested that tumor uptake of HPMA 

copolymers may vary from tumor to tumor and with different tumor sizes. It might affect the 

efficacy of the polymer conjugates and more detailed studies are needed to understand the 

possible correlations. 

FIGURE 11.2 Three-dimensional maximum intensity projection (MIP) (a) and 2D coronal MR images (b) of 

the male nude mouse bearing prostate carcinoma DU-145 (left) and human Kaposi’s sarcoma (SLK) tumor 

(right) at different time points. The arrow points to the tumor. Imaging parameters were TRZ7.8 ms, 

TEZ2.7 ms, 258 flip angle, and 0.4-mm coronal slice thickness. 


206 

11.3.2 MR IMAGING OF A PARAMAGNETIC HPMA CONJUGATE FOR BONE DELIVERY 

Kopecek and coworkers labeled HPMA copolymers (55 kDa) with Gd-DOTA to noninvasively 

evaluate macromolecular accumulation in arthritic joints in a rat model. The copolymers were 

labeled with fluorescein in addition to the paramagnetic metal chelate. The structure of the conjugate 

is shown in Figure 11.3. After intravenous administration of the conjugate in a rat model with 

adjuvant-induced arthritis, the biodistribution of copolymers was followed by contrast-enhanced 

MRI at various time points post-injection. 35 

Contrast-enhanced dynamic MRI clearly revealed the pharmacokinetic properties and biodistribution 

of the conjugate with high anatomic resolution in the adjuvant-induced arthritic rats 

(Figure 11.4). The copolymers had a relatively high molecular weight and a considerable 

amount of the copolymers was still circulating in the blood 3 h post-injection. Selective accumulation 

of HPMA copolymers was gradually shown in the arthritic joints of the adjuvant-induced 

arthritis rats. The copolymers were also cleared from the accumulation sites over time as shown by 

contrast-enhanced MRI. The uptake and retention of the MR contrast agent labeled polymer 

correlated well with the histopathological features of inflammation and local tissue damages. 

No significant contrast enhancement was observed in the hind-limb joints of healthy rats in 

a control study. 

11.4 MR IMAGING OF PARAMAGNETIC POLY(L-GLUTAMIC ACID) 

CONJUGATES 

Poly(L-glutamic acid) (PGA) is a negatively charged, biocompatible polymer that has been used as 

a carrier for drug delivery. 36 It is a homopolymer of L-glutamic acid linked through peptide bonds 

and the pendant g-carboxyl groups are suitable to load therapeutic agents via chemical conjugation. 

Several poly(L-glutamic acid)-anti-cancer drug conjugates have been developed for cancer treatment. 

37–39 The conjugation of a MRI contrast agent to PGA enables noninvasive investigation of its 

pharmacokinetics and in vivo drug delivery with contrast-enhanced MRI. 

11.4.1 MR IMAGING OF PGA–(GD-DOTA) CONJUGATE IN AN ANIMAL TUMOR MODEL 

Poly(L-glutamic acid)–1,6-hexanediamine–(Gd-DOTA) conjugate was prepared using a synthetic 

procedure similar to that of PGA-cystamine-Gd-DOTA conjugate synthesis 40 to noninvasively 

HO 

CH 2 

S 

CH3 C 

m 

O 

NH 

CH2 CH2 CH2 NH 

NH 

O 

CH 2 

COOH 

O 

CH3 C 

n 

O 

NH 

CH2 CH OH 

CH3 CH 2 

O 

- OOC 


CH3 C 

O 

NH 

CH2 CH 

CH2 NH 

N 

Gd 3+ 

N 

N N 

COO - 

COO - 

FIGURE 11.3 The structure of HPMA-MAFITC copolymers labeled with Gd(III)-DOTA. 



FIGURE 11.4 The MR images of rats taken at different time points. The acquired images were post-processed 

using the maximum intensity projection (MIP) algorithm. P-Gd(III)-DOTA is abbreviated as P-Gd. (A-H) AIA 

rat images at baseline (a) and 5 min (b), 1 h (c), 2 h (d), 3 h (e), 8 h (f), 32 h (g), 43 h (h) post-injection of 

P-Gd(III)-DOTA. (i–n) Healthy rat images at baseline (i) and 5 min (j), 1 h (k), 2 h (l), 8 h (m), 48 h (n) 

post-injection of P-Gd(III)-DOTA. Arrow points to the diseased joint (Adapted from Wang, D. etal., Pharm. 

Res., 21, 1741, 2004. With permission.) 

study delivery efficiency of the conjugate. The structure of the conjugate is shown in Figure 11.5. 

PGA–1,6-hexanediamine–(Gd-DOTA) conjugates, H a high molecular weight (50 kDa, PDZ 

1.12) conjugate and a low molecular weight conjugate (20 kDa, PDZ1.08) were investigated by 

contrast-enhanced MRI to study the size effect of conjugates on the efficiency of in vivo drug 

delivery. The conjugates with different molecular weights had similar Gd content (41.7% for 

50 kDa conjugate and 40.2 mol% for 20 kDa conjugate) and T1 relaxivity (9.44 and 9.20 mM K 

1 s K1 for 50 and 20 kDa conjugates, respectively). The conjugates were intravenously administered 

into female nu/nu athymic mice bearing human breast carcinoma MB-231 xenografts at a dose of 

0.07 mmol-Gd/kg. Dynamic contrast-enhanced MRI clearly revealed the size effect of the conjugates 

on their pharmacokinetics and in vivo tumor delivery efficiency in the animal model. 

Figure 11.6 shows the T1-weighted coronal MR images of mice before and at various time 

points after injection with the conjugates. Strong contrast enhancement was observed in the heart at 

the initial stage post-injection and then gradually faded away for both conjugates. The signal 

intensity decreased more rapidly for the low molecular weight conjugate than the high molecular 

weight conjugate, validating the concept that the blood circulation of the conjugates is prolonged 

NH 

CH 

CH2 CH2 C O 

OH 

O 

C 

m 

NH 

CH 

CH2 CH2 C O 

O 

C 

n 

NH(CH 2) 6NH 

O 

C 

− OOC 

N N 

Gd 

N N 

3+ 

COO − 

COO − 

FIGURE 11.5 The structure of poly(L-glutamic acid)–1,6-hexanediamine–(Gd-DOTA) conjugate. 


208 

FIGURE 11.6 Coronal MR slice images of mice bearing MB-231 breast cancer tumors using PGA–1,6-hexanediamine–(Gd-DOTA) 

conjugates with different molecular weights. The images were taken at different time 

points using a 3D FLASH pulse sequence (1.74-ms TE, 4.3-ms TR, 258 RF tip angle, 120-mm FOV, and 

1.6-mm coronal slice thickness). 

with increase in molecular weights. The low-molecular-weight conjugate (20 kDa) was cleared 

from the blood circulation within 30 min after-injection. The dynamic changes in signal intensity in 

the liver correlated well to that in the heart. The enhancement in the liver returned to the background 

level 24 h after the injection, indicating low liver uptake of the PGA conjugates. 

The contrast enhancement pattern in the tumor tissue was similar to that of HPMA conjugates 

in Kaposi’s sarcoma. Strong signal was observed at tumor periphery and little in the inner tumor 

tissue. The dynamic MR images showed that the intensity and duration of tumor enhancement 

strongly depended on the size of the conjugates, indicating size-dependent tumor accumulation. 

Rapid clearance of the low molecular weight conjugate (20 kDa) from the blood circulation 

resulted in low and short accumulation in the tumor tissue. The high molecular weight conjugate 

(50 kDa) had a long blood circulation time, resulting in an increased contrast enhancement for at 

least 4 h. The enhancement also gradually extended into inner tumor tissue, indicating diffusion of 

the conjugate further into the tumor. The results clearly showed that drug delivery with polymeric 

conjugate into tumor tissue was a slow process and H conjugates with high molecular weight and 

long blood circulation time was more effective for tumor delivery than the conjugate with low 

molecular weight. Contrast-enhanced MRI clearly revealed the size effect of the conjugates on their 

pharmacokinetics and in vivo drug delivery efficiency. 

11.4.2 MR IMAGING OF PGA–MCE6–(Gd-DOTA) CONJUGATE 

IN AN ANIMAL TUMOR MODEL 


Most anti-cancer drugs are lipophilic and the conjugation of lipophilic drugs to a polymeric carrier may 

also modify the in vivo behavior of the polymers. Mesochlorin e 6 (Mce 6) is a lipophilic photosentizer 

for photodynamic therapy. 41,42 Aparamagneticpoly(L-glutamic acid) Mce6 conjugate was prepared to 

investigate the pharmacokinetics, biodistribution and drug delivery efficiency of the PGA–Mce6 conjugate 

with contrast-enhanced MRI. 43 The structure of the conjugate is shown in Figure 11.7. A PGA–1,6hexanediamine–(Gd-DOTA) 

conjugate was also prepared from the same poly(L-glutamic acid) as a 

control. Mce6 content was 2.5 mol% in PGA–Mce6–1,6-hexanediamine–(Gd-DOTA) and Gd content 

was 20 and 29 mol% for PGA–Mce6–1,6-hexanediamine–(Gd-DOTA) and PGA–1,6-hexanediamine– 

(Gd-DOTA), respectively. The T 1 relaxivity was 8.46 and 8.33 mM K1 s K1 for PGA–Mce 6–1,6-hexanediamine–(Gd-DOTA) 

and PGA–1,6-hexanediamine–(Gd-DOTA), respectively. The small amount 



NH 

CH 

CH2 CH = 2 

C O 

NH 

CH2 O 

CH2 NH 

O 

C 

m 

HOOC 

NH 

CH 

CH2 CH2 C O 

OH 

NH N 

mesoc ( 

HOOC 

N 

HN 

O 

C 

n 

− OOC 

N N 

Gd 

N N 

3+ 

COO − 

COO − 

of Mce6 in PGA–Mce6–1,6-hexanediamine–(Gd-DOTA) significantly reduced the hydrodynamic 

volume and molecular weight distribution of the conjugate. The weight (Mw) and number (Mn) 

average molecular weights decreased from 101 and 49 kDa of PGA–1,6-hexadiamine–(Gd-DOTA) 

to 49 and 34 KDa for PGA–Mce6–1,6-hexadiamine–(Gd-DOTA). The lipophilic Mce6 may alter the 

conformation of the polymers, resulting in a smaller hydrodynamic volume for PGA–Mce6–1,6-hexadiamine–(Gd-DOTA). 

A contrast-enhanced MRI study demonstrated that PGA–Mce6–1,6-hexanediamine–(Gd-DOTA) 

had significantly different pharmacokinetics and biodistribution from PGA–1,6-hexanediamine–(Gd- 

DOTA) in female nu/nu athymic mice bearing human breast carcinoma MB-231 xenografts. 

Figure 11.8 shows coronal dynamic MR images of mice injected with PGA–Mce6–1,6-hexanediamine–(Gd-DOTA) 

and PGA–1,6-hexanediamine–(Gd-DOTA) at a dose of 0.07 mmol-Gd/kg. 

The conjugates demonstrated different dynamic contrast enhancement patterns in the heart and liver. 

PGA–Mce 6–1,6-hexadiamine–(Gd-DOTA) had a strong contrast enhancement in the liver and 

relatively weak enhancement in the heart, whereas PGA–1,6-hexadiamine–(Gd-DOTA) exhibited 

NH 

CH 

CH2 CH2 C O 

O 

C 

NH(CH 2) 6NH 

FIGURE 11.7 The structure of PGA–Mce6–1,6-hexanediamine–(Gd-DOTA). 

FIGURE 11.8 Coronal contrast-enhanced MR slice images for PGA–Mce 6–(Gd-DOTA) conjugate and 

PGA–(Gd-DOTA) conjugate through the heart (b) before and at 5, 15, 30, 60, and 120 min post-injection. 


l 

O 

C

210 

strong enhancement in the heart and relatively weak enhancement in the liver. The conjugation of 

Mce 6 significantly altered the pharmacokinetics and biodistribution of the poly(L-glutamic acid). The 

results indicated that PGA–Mce 6–1,6-hexadiamine–(Gd-DOTA) had higher liver accumulation than 

PGA–1,6-hexadiamine–(Gd-DOTA). Consequently, it cleared from the blood circulation more rapidly 

than PGA–1,6-hexadiamine–(Gd-DOTA), which could also be attributed to the larger 

hydrodynamic volume of PGA–1,6-hexadiamine–(Gd-DOTA). 

11.5 SUMMARY 

Contrast-enhanced MRI is effective for noninvasive visualization of the real-time pharmacokinetics 

and biodistribution of paramagnetically labeled polymer drug conjugates after intravenous administration. 

The technique provides three-dimensional anatomic images of soft tissues with high 

spatial resolution. The circulation and accumulation of the paramagnetic polymeric conjugates 

in organs or tissues result in bright signal for T 1-weighted images. Dynamic contrast-enhanced 

MRI clearly reveals circulation of the conjugates in the blood, interaction with the major organs, 

including the liver, heart, kidneys, etc., and accumulation and clearance in the target tissues. It also 

provides information on the dynamic changes of the distribution patterns of polymer conjugates in 

solid tumor tissues, which cannot be obtained with conventional pharmacokinetic methods. Such a 

detailed map of the delivery system deposition within the tissue and its correlation with the local 

pathologic features may help to optimize the structure of the polymeric drug conjugate to achieve 

high drug delivery efficiency. One limitation of contrast-enhanced MRI is the accurate quantification 

of the concentrations of the conjugates in the tissues and organs. Several technical factors 

control the accurate measurement of concentrations of contrast agents with conventional contrastenhanced 

MRI. For quantitative study of in vivo drug delivery with paramagnetic conjugates, MR 

T1 mapping can be used because T1 relaxation rate (1/T1) is linearly proportional to the concentration 

of MRI contrast agents. Data acquisition and algorithms to process the data for T1 mapping 

are more complicated than for conventional contrast-enhanced MRI. 

REFERENCES 


1. Kopecek, J. et al., HPMA copolymer-anticancer drug conjugates: design, activity and mechanism of 

action, Eur. J. Pharm. Biopharm., 50, 61, 2000. 

2. Kopecek, J. et al., Water soluble polymers in tumor targeted delivery, J. Control. Rel., 74, 147, 2001. 

3. Hobbs, S. K. et al., Regulation of transport pathways in tumor vessels: role of tumor type and 

microenvironment, Proc. Natl Acad. Sci. USA, 95, 4607, 1998. 

4. Maeda, H., Sawa, T., and Konno, T., Mechanism of tumor-targeted delivery of macromolecular drugs, 

including the EPR effect in solid tumor and clinical overview of the prototype polymeric drug 

SMANCS, J. Control. Rel., 74, 47, 2001. 

5. Minko, T., Kopeckova, P., and Kopecek, J., Chronic exposure to HPMA copolymer-bound adriamycin 

does not induce multidrug resistance in a human ovarian carcinoma cell line, J. Controlled Rel., 

59, 133, 1999. 

6. Minko, T. et al., HPMA copolymer bound adriamycin overcomes MDR1 gene encoded resistance in a 

human ovarian carcinoma cell line, J. Control. Rel., 54, 223, 1998. 

7. Abe, S. and Otsuki, M., Styrene maleic acid neocarzinostatin treatment for hepatocellular carcinoma, 

Curr. Med. Chem. Anti-Cancer Agents, 2, 715, 2002. 

8. Brewerton, L. J., Fung, E., and Snyder, F. F., Polyethylene glycol-conjugated adenosine phosphorylase: 

development of alternative enzyme therapy for adenosine deaminase deficiency, Biochim. 

Biophys. Acta, 1637, 171, 2003. 

9. Pinheiro, J. P. et al., Drug monitoring of PEG-asparaginase treatment in childhood acute lymphoblastic 

leukemia and non-Hodgkin’s lymphoma, Leuk. Lymphoma, 43, 1911, 2002. 

10. Ettinger, L. J. et al., An open-label, multicenter study of polyethylene glycol-L-asparaginase for the 

treatment of acute lymphoblastic leukemia, Cancer, 75, 1176, 1995. 



11. Motzer, R. J. et al., Phase II trial of branched peginterferon-alpha 2a (40 kDa) for patients with 

advanced renal cell carcinoma, Ann. Oncol., 13, 1799, 2002. 

12. Rowinsky, E. K. et al., A phase I and pharmacokinetic study of pegylated camptothecin as a 1-hour 

infusion every 3 weeks in patients with advanced solid malignancies, J. Clin. Oncol., 21, 148, 2003. 

13. Auzenne, E. et al., Superior therapeutic profile of poly-L-glutamic acid-paclitaxel copolymer 

compared with Taxol in xenogeneic compartmental models of human ovarian carcinoma, Clin. 

Cancer Res., 8, 573, 2002. 

14. Gianasi, E. et al., HPMA copolymer platinates as novel antitumor agents: in vitro properties, pharmacokinetics 

and antitumor activity in vivo, Eur. J. Cancer, 35, 994, 1999. 

15. Thomson, A. H. et al., Population pharmacokinetics in phase I drug development: a phase I study of 

PK1 in patients with solid tumors, Br. J. Cancer, 81, 99, 1999. 

16. Seymour, L. W. et al., Hepatic drug targeting: phase I evaluation of polymer-bound doxorubicin, 

J. Clin. Oncol., 20, 1668, 2002. 

17. Liang, Z. P. and Lauterbur, P. C., Principles of Magnetic Resonance Imaging, IEEE Press, New York, 

1999. 

18. Vlaardingerbroek, M. T. and den Boer, J. A., Magnetic Resonance Imaging, 3rd ed., Springer, New 

York, 2003. 

19. Lauffer, R. B., Paramagnetic metal complexes as water proton relaxation agents for NMR imaging: 

theory and design, Chem. Rev., 87, 901, 1987. 

20. Merbach, A. E. and Toth, E., The Chemistry of Contrast Agents in Medical Magnetic Resonance 

Imaging, John Wiley & Sons, Inc., England, 2001 chap. 1–2 

21. Weinmann, H. J. et al., Characteristics of gadolinium-DTPA complex: a potential NMR contrast 

agent, Am. J. Roentgenol., 142, 619, 1984. 

22. Kaplan, G. D., Aisen, A. M., Aravapalli, S. R. et al., Preliminary clinical trial of gadodiamide 

injection: a new nonionic gadolinium contrast agent for MR imaging, J. Magn. Reson. Imaging, 1, 

57, 1991. 

23. Magerstadt, M. et al., Gd(DOTA): An alternative to Gd(DTPA) as a T1,2 relaxation agent for NMR 

imaging or spectroscopy, Magn. Reson. Med., 3, 808, 1986. 

24. Tweedle, M. F., The proHance story: the making of a novel MRI contrast agent, Eur. Radiol., 

7(Suppl 5), 225, 1997. 

25. Youk, J. H., Lee, J. M., and Kim, C. S., MRI for detection of hepatocellular carcinoma: comparison of 

mangafodipir trisodium and gadopentetate dimeglumine contrast agents, Am. J. Roentgenol., 183, 

1049, 2004. 

26. Weissleder, R. et al., Ultrasmall superparamagnetic iron oxide: an intravenous contrast agent for 

assessing lymph nodes with MR imaging, Radiology, 175, 494, 1990. 

27. Wang, S. C. et al., Evaluation of Gd-DTPA-labeled Dextran as an intravascular MR contrast agent: 

imaging characteristics in normal rat tissues, Radiology, 175, 483, 1990. 

28. Bryant, L. H. et al., Synthesis and relaxometry of high-generation (GZ5, 7, 9 and 10) PAMAM 

dendrimer-DOTA-gadolinium chelates, J. Magn. Reson. Imaging, 9, 348, 1999. 

29. Lu, Z. R. et al., Design of novel bioconjugates for targeted drug delivery, J. Control. Rel., 78, 165, 

2002. 


31. Pimm, M. V. et al., Gamma scintigraphy of the biodistribution of 123 I labeled N-(2-hydroxypropyl)methacrylamide 

copolymer-doxorubicin conjugates in mice with transplanted melanoma and 

mammary carcinoma, J. Drug Target, 3, 375, 1996. 

32. Kissel, M. et al., Synthetic macromolecular drug carriers: biodistribution of poly [(N-2-hydroxypropyl)methacrylamide] 

copolymers and their accumulation in solid rat tumors, J. Pharm. Sci. Tech., 55, 

191, 2001. 

33. Mitra, A. et al., Technetium-99m-Labeled N-(2-hydroxypropyl) methacrylamide copolymers: 

synthesis, characterization, and in vivo biodistribution, Pharm. Res., 21, 1153, 2004. 

34. Wang, Y.L., and Lu, Z.R. Unpublished data 2005. 

35. Wang, D. et al., The arhrotropism of macromolecules in adjuvant-induced arthritis rat model: a 

preliminary study, Pharm. Res., 21, 1741, 2004. 

36. Li, C., Poly (L-glutamic acid)-anticancer drug conjugates, Adv. Drug Deliv. Rev., 54, 695, 2002. 


212 


37. Hurwitz, E., Wilchek, M., and Pitha, J., Soluble macromolecules as carriers for daunorubicin, J. Appl. 

Biochem., 2, 25, 1980. 

38. Li, C. et al., Complete regression of well-established tumors using novel water-soluble poly (Lglutamic 

acid)-paclitaxel conjugates, Cancer Res., 58, 2404, 1998. 

39. Zou, Y. et al., Effectiveness of water soluble poly(L-glutamic acid)–camptothecin conjugate against 

resistant human lung cancer xenografted in nude mice, Int. J. Oncol., 18, 331, 2001. 

40. Lu, Z. R. et al., Poly(L-glutamic acid) Gd(III)-DOTA conjugate with a degradable spacer for magnetic 

resonance imaging, Bioconjugate Chem., 14, 715, 2003. 

41. Lu, Z. R., Kopeckova, P., and Kopecek, J., Polymerizable Fab’ fragment for targeting of anticancer 

drugs, Nature Biotechnol., 17, 1101, 1999. 

42. Lu, Z. R. et al., Polymerizable Fab’ antibody fragments targeted photodynamic cancer therapy in nude 

mice, STP Pharm. Sci., 13, 69, 2003. 

43. Vaidya, A. et al. Non-invasive in vivo imaging of a polymeric drug conjugate using contrast enhanced 

MRI. In Transactions of 32nd Annual Meeting of the Controlled Release Society, 2005. 


12 

CONTENTS 

Polymeric Nanoparticles for 

Tumor-Targeted Drug Delivery 

Tania Betancourt, Amber Doiron, 

and Lisa Brannon-Peppas 

12.1 Introduction ........................................................................................................................215 

12.2 Targeting to Cancer............................................................................................................216 

12.3 Passive Targeting and the EPR Effect ...............................................................................217 

12.4 Targeting to Angiogenesis .................................................................................................218 

12.4.1 Targeting Using Vascular Endothelial Growth Factor Receptors...................... 218 

12.4.2 Targeting Using Integrins ................................................................................... 218 

12.4.2.1 Integrins as Targets for Imaging .........................................................220 

12.5 Targeting Using Folate Receptors .....................................................................................220 

12.5.1 Antibodies and Folate Receptors ........................................................................ 221 

12.5.2 Folate-Targeted Nanoparticles for Gene Delivery ............................................. 222 

12.6 Approaches for Cancer Targeting to Specific Cancer Types ............................................223 

12.6.1 Prostate Cancer.................................................................................................... 224 

12.7 Targeted Nanoparticles and Imaging of Cancer................................................................225 

12.8 Other Targets for Cancer ...................................................................................................225 

12.9 Avidin and Biotin Targeting ..............................................................................................226 

12.10 Conclusions ........................................................................................................................226 

References.....................................................................................................................................226 


Cancer is a disease that affects millions of people across the globe every year. The World Health 

Organization estimated that more than 10 million people developed a malignant tumor and more 

than 6.5 million people died from this disease during the year 2000. 1 In the United States, cancer is 

the second cause of deaths from disease after heart disease, accounting for more than half a million 

deaths every year. According to the American Cancer Society (ACS) cancer statistics, the overall 

cost for cancer for the United States in 2004 was $189.8 billion: $69.4 billion for direct medical 

costs, $16.9 billion for indirect morbidity costs, and $103.5 billion for indirect mortality costs. 2 

Furthermore, while mortality rates of other major chronic diseases, such as heart and cerebrovascular 

disease, decreased significantly in the past half-century, cancer mortality rates have 

remained approximately constant. 2 This is a troubling fact because it suggests that recent detection 

and treatment options have not been able to improve mortality rates substantially. 


215

216 

Research in the past decade has focused on using unique characteristics of cancer cells and the 

vasculature surrounding those cells to deliver imaging agents, chemotherapeutic drugs, gene 

therapy, and other active agents directly and selectively to cancerous tissues. Many of these new 

formulations (described elsewhere in this volume) are liposomes, prodrugs, polymer conjugates, 

micelles, and dendritic systems. This chapter will concentrate on polymeric nanoparticles that have 

been studied as targeted systems for treatment and detection of cancer. 

12.2 TARGETING TO CANCER 


Nanoparticles may be targeted to the growing vasculature serving the growing cancer or to the 

cancer cells themselves. Targeted delivery utilizes unique phenotypic features of diseased tissues 

and cells in order to concentrate the drug at the location where it is needed. Targeted delivery can be 

divided into passive and active targeting. Passive targeting tries to minimize nonspecific 

interactions between the drug carrier and nontarget sites in the body by detailing the physiochemical 

properties of the aberrant tissue such as size, morphology, hydrophilicity, and surface charge. 3 

When targeting tumor tissue, the enhanced permeability and retention effect (EPR) is an example of 

passive targeting approach; it allows passage of drug carriers ranging in size from 10 to 500 nm 

through the highly permeable blood vessels that supply growing tumors and leads to entrapment of 

large molecules as a result of deficient lymphatic drainage. 3–5 In fact, it has been reported that the 

intra-cellular openings in vascular endothelium of tumor blood vessels can be up to 2 mm in 

diameter and that the vessel leakiness in tumor vasculature can be up to an order of magnitude 

higher than that of normal blood vessels. 5 Active targeting utilizes biologically specific interactions 

including antigen–antibody and ligand–receptor binding and may seek drug uptake by receptormediated 

endocytosis through association of the drug or drug carrier with such antigen or ligand. 3 

Receptor-mediated endocytosis commonly occurs through clathrin-coated vesicles and is carried 

out in mammalian cells continuously for the uptake of nutrients and for modulation of signal 

transduction through the up- or down-regulation of signaling receptors. 6 Targeted delivery 

avoids the need for high systemic drug levels for the drug to be effective and consequently 

offers a more economic alternative for treatment. Not only is it useful for therapeutic purposes; 

it is also beneficial in diagnosis. Recent research has pointed to its ability to concentrate imaging or 

contrast agents for the detection of malignancies and for monitoring the effects of therapeutic 

agents. 7,8 To date, most systems for targeted delivery have utilized drug conjugates, liposomes 

or micelles. 9–11 Targeting of particulate systems has focused more often on passive targeting based 

on size than on active targeting. But systems that combine both methods, starting with passive 

targeting through EPR and enhancing the targeting through specific interactions are beginning to 

show great promise. 

While it is challenging to deliver a drug or imaging agent-containing nanoparticle directly and 

selectively to a cancerous cell or tissue, the additional challenges of then having that particle and/or 

its contents being transported into the targeted cell have often been overlooked. Couvreur presented 

some of these challenges at the 11th International Symposium on Recent Advances in Drug 

Delivery Systems and also summarized that presentation in a recent publication. 12 The tumor 

resistance can be due to the deliverance of nanoparticles to the tumor as well as to resistance to 

the active agent being delivered. In this article, many different pathways are described and the 

enhancement of drug delivery due to the presence of nanoparticles, especially polycyanoacrylate 

nanoparticles, is evaluated and summarized. The enhancement of drug permeability into cells due 

to interactions with biodegradation byproducts as well as the effect of nanoparticle surface charge 

is discussed. 

In this chapter, we will first review recent work on targeting to the two most promising targets 

for cancer: Angiogenesis and folate receptors. We will then describe other potential targets for 

cancer imaging and therapy with nanoparticles, including antibodies strategies using biotin. 


Polymeric Nanoparticles for Tumor-Targeted Drug Delivery 217 

12.3 PASSIVE TARGETING AND THE EPR EFFECT 

Although active targeting may be achieved through targeting to folate receptors, angiogenesis, or 

targets more specific to the cancer being treated, some enhancement of treatment to cancers on 

account of the EPR effect cannot be ignored and should be utilized until more specific targeting 

systems can be developed. Preparation of nanoparticle systems that can avoid uptake by the 

reticuloendothelial system (RES) is essential, and such particles are often referred to as 

“stealth” nanoparticles. 

The most effective polymer used as a coating on nanoparticles to avoid detection by the RES is 

poly(ethylene glycol). 13 The latter can be achieved by adsorption of the PEG-containing polymer 

onto the surface of nanoparticles, direct conjugation of the PEG to the nanoparticles, or inclusion of 

PEG in the polymeric backbone that makes up the nanoparticles. The PEG itself may then also be 

modified to give targeting capabilities and to avoid uptake by the RES. The PEG works to mask the 

surface of the nanoparticles by reducing the plasma and protein adsorption to the particles, reducing 

the complement activation and hence the recognition of the PEG as a foreign substance in the blood 

stream. The longer circulation time afforded PEGylated nanoparticles allows them time to be 

targeted, whether passively or actively, to cancerous tissues. 

Extended circulation time and enhanced tumor targeting were seen for poly(ethylene oxide)modified 

poly(epsilon-caprolactone) nanoparticles in mice with tumors of MDA-MB-231, a human 

breast carcinoma. 14 These particles were loaded with [ 3 H]-tamoxifen. The amount of labeled 

tamoxifen at the tumor site, 6 h after injection, for those particles with PEO modification was at 

least twice that of particles without modification and four times that of labeled tamoxifen injection. 

The amount of labeled tamoxifen found in the blood stream at 6 h after injection was also at least 

twice that of the injection or unmodified nanoparticle formulations. 

The stability and circulation of PLGA–mPEG nanoparticles containing cisplatin was investigated. 

It was found that while the mPEG content affected the drug release rate, the drug loading 

level had no effect on the drug release rate for in vitro studies. 15 The release was more that 60% 

completed within the first 12 h in all cases. Data for blood levels was only presented for 3 h, so it is 

hard to draw any valid conclusions from this information. 

Nanoparticles of PLA and PEG–PPG–PEG were prepared containing irinotecan, a prodrug of 

an analogue of camptothecin. 16 Although little characterization beyond the average particles size 

(231 nm) was presented here and no in vitro studies were described, the in vivo studies are quite 

interesting. In this study, there was a modest increase in survival time in mice with M5076 tumors 

(early liver metastatic stage) after a single injection and more pronounced increases in survival time 

after either two or three repeat injections. It is noteworthy that the greatest survival times (20% 

survival at 45 days at the end of the study) were seen with two injections at days three and five after 

the implantation of the tumor. 

Polycyanoacrylate nanoparticles have long been studied by Couvreur and collaborators as 

biodegradable nanoparticles for a variety of applications and therapies which are not limited to 

treatment of cancer. 17 A recent work describes the effectiveness of these nanoparticles as delivery 

systems for brain tumor targeting. Here they studied uncoated and PEG-coated nanoparticles and 

found that both types of particles showed accumulation in a well-established 9L gliosarcoma in rat 

studies. The PEG-coated particles showed the highest accumulation, with a tumor-to-brain ratio 

of 11. 

Poly(butyl cyanoacrylate) nanoparticles containing doxorubicin were prepared with no surface 

modifications; the in vivo distribution of 99m Tc labeled nanoparticles was evaluated in mice inoculated 

with Dalton’s lymphoma tumor cells. 18 The nanoparticles were administered by subcutaneous 

injection and the concentration in a number of organs was followed for 48 hours and compared with 

that for 99m Tc labeled doxorubicin alone. When compared with the amount of doxorubicin alone, 

the amount of radioactivity in the tumor was higher at all times tested, with a 13-fold increase seen 

at 48 h. 


218 

While the majority of work on targeted nanoparticles has been carried out with polylactides, 

polyglycolides and polycyanoacrylates, those are not the only materials that can be used in nanoparticles. 

Some recent work with radiolabeled gelatin nanoparticles modified with poly(ethylene 

glycol) showed that adding the PEG to the surface of these nanoparticles increased the circulation 

time in mice with double the amount of PEG-modified nanoparticles in the blood stream 3 h after 

injection compared to the amount of control nanoparticles. 19 In addition, there was a four-fold 

increase in the amount of PEG-modified gelatin nanoparticles found in the tumor 4 h after injection 

and later when compared to the number of nonmodified gelatin nanoparticles. 

12.4 TARGETING TO ANGIOGENESIS 

A recently explored and potentially promising target of cancer drug and gene nanoparticle therapy 

is tumor angiogenesis. It is now well established that tumor growth is dependent on new capillary 

infiltration from surrounding, preexisting vasculature. 20,21 This is an important control point in 

cancer as much research has proven that tumors cannot effectively grow past a small size or 

metastasize without blood supply. 22–24 Except for the cases of menstruation, wound healing, and 

tissue regeneration, capillaries do not increase in size or number under normal physiological 

conditions. Tumor growth is an exception to this physiological rule. 

Tumors are typically unable to affect angiogenesis when they are small and surrounded by 

healthy tissue. However, at the point in growth where nutrients, oxygen, and growth factors can no 

longer reach the cancer cells, blood flow is required to allow further growth of the tumor. After what 

is occasionally a substantial time period, the tumor may abruptly induce angiogenesis into the 

tissue. 23 Because understanding this step in the progression of cancer is thought to be of great 

importance, much research has been and continues to be focused on pinpointing the progression of 

cancer and on targeting the event for therapy. 

Much research has focused on targeted therapy of either chemotherapeutic agents to sites of 

tumor angiogenesis or of angiogenesis-inhibiting drugs to tumors with the goal of directly combatting 

the proliferation of newly forming capillaries in the tumor. Angiogenesis is a complex, multicomponent 

process that involves many cell types, cytokines, growth factors and receptors, 

proteases, and adhesion molecules. 25 As a result, there are many potential targets for anti-angiogenic 

or chemotherapeutic therapy. Some recent advances in targeting approaches for nanoparticle drug 

delivery are discussed below. 

12.4.1 TARGETING USING VASCULAR ENDOTHELIAL GROWTH FACTOR RECEPTORS 

Vascular endothelial growth factor (VEGF) is particularly important in the process of angiogenesis 

and has been shown to greatly affect tumor growth in animal models. 26–28 The VEGF receptors have 

been used as a means to target the vascular bed in many instances. VEGF receptor-2 (VEGFR-2) 

has recently been used to target nanoparticles to tumor vascular beds by Li et al in mice with 

K1735-M2 tumors. 29 A succinyl-dextran-polymerized nanoparticle conjugated to rat anti-mouse 

VEGFR-2 antibody and radioisotope 90 Y caused a significant tumor growth delay compared to 

conventional radiolabeled antibody and other controls. Additionally, anti-CD31 staining showed 

a decreased vessel density and damage to tumor vessels after treatment with the anti-VEGFR- 

2- 90 Y nanoparticles. 

12.4.2 TARGETING USING INTEGRINS 


The integrins represent another important cell surface molecule group for angiogenesis targeting 

because some integrins, such as avb3 and avb5, are upregulated on the endothelial cell surface of 

neovasculature. 30,31 The a vb 3 integrin is expressed on numerous tumor cell types; it is highly 

expressed on neovascular endothelial cells. Hood and collaborators used 40-nm diameter 



cationic-lipid-based nanoparticles coupled to an organic avb3 ligand that was shown to be specific 

for avb3 in cell studies. These nanoparticles, which contained a luciferase reporter plasmid, were 

injected into mice with the avb3-negative cell line M21-L. The nanoparticles targeted to the 

neovasculature within the tumor (but not to the tumor cells themselves) with no expression elsewhere 

in the mice as detected by luciferase expression. In order to test therapeutic efficacy, NPs 

were conjugated to a mutant Raf gene that blocks angiogenesis. Systemic injection in the M21-L 

tumor-expressing mice rapidly induced apoptosis of endothelial cells within the tumor. Tumor 

regression was seen within 10 days. 32 

In the same study with VEGFR-2 targeting discussed above, Li and collaborators also demonstrated 

targeting of the radioisotope 90 Y using nanoparticles targeted to the integrin avb3 with a 

small molecule integrin agonist. 29 In mice with K1735-M2 tumors, avb3-targeted 90 Y-nanoparticles 

significantly delayed tumor growth compared to untreated tumors. TUNEL staining of tumor 

sections showed widespread apoptosis in tumors treated with these targeted nanoparticles. The 

authors have postulated that this targeted nanoparticle radiotherapy has the potential to be used 

to treat a variety of solid tumors. They have also postulated that the use of nanoparticles increases 

efficacy due to the high payload delivered by the carriers. 

Additionally, PEGylated polyethyleneimine (PEI) nanoplexes with a cyclic disulfide bond 

constrained Arg–Gly–Asp (RGD) peptide ligand at the distal end have been used to target 

integrin-expressing tumor neovasculature to deliver siRNA. Integrins are receptors for extracellular 

matrix components that contain a tripeptide RGD sequence. Therefore, RGD containing peptide 

sequences can target to cell surface integrins that are upregulated on neovasculature. The siRNA 

used inhibited angiogenesis by inhibiting VEGFR-2 expression. Intravenous injection of nanoparticles 

into nude mice with N2A tumors showed tumor uptake of the siRNA, inhibition of protein 

synthesis in the tumor, and inhibition of angiogenesis and tumor growth. 33 This study demonstrates 

tumor selective delivery through both the targeting ligand and gene pathway by using siRNA. 

In another a vb 3 targeting approach using an RGD peptide, Kopelman has created multifunctional 

nanoparticles of 30–60 nm for the treatment of gliomas. 34 The nanoparticles are able to kill 

cancer cells by bombarding them with externally released reactive oxygen species created by 

photodynamic agents activated by laser light. The particles also contain superparamagnetic iron 

oxide and enhance imaging by magnetic resonance. The photodynamic sensitizer and MRI contrast 

agents are entrapped within a polyacrylamide core, the surface of which is coated with PEG chains 

and targeting RGD moieties. The particles containing photodynamic agents were shown to produce 

sufficient singlet oxygen to kill cells in vitro. Additionally, these nanoplatforms were injected into 

an in vivo rat intracerebral 9L tumor model, and diffusion MRI was performed at various times to 

evaluate the tumor diffusion, tumor growth, and tumor load. The gliomas treated with the nanoparticles 

and irradiated with laser light caused regional necrosis and significant shrinkage of tumor 

mass, a shrinkage that lasted for 12 days. The authors postulate that the light activated release of 

reactive oxygen from photosensitizer-containing nanoparticles is a viable approach for brain tumor 

treatment. Also, the incorporation of MRI contrast agents allows for monitoring of treatment and 

tumor progression in vivo. 

Carbohydrate based nanoparticles have also been used to target drugs to neovasculature via an 

avb3/cyclic RGD peptide interaction. Inulin multi-methacrylate formed the core of the nanoparticles 

and was attached to the RGD targeting moiety with a PEG linker. Doxorubicin was loaded in 

the nanoparticles via covalent and noncovalent linkages. The pharmacokinetics and biodistribution 

of the doxorubicin loaded nanoparticles were studied over five days in female Balb/cJ mice with 

metastatic mammary tumor clone-66. A bi-exponential fix with a terminal half-life of 5.99 h was 

observed; decreasing drug concentrations with time in the heart, lungs, kidney, and plasma was also 

observed. Conversely, increasing drug accumulation was observed in the liver, spleen, and in the 

tumor where there was also the presence of high levels of doxorubicin metabolite. The presence of 

the high metabolite levels in the tumorsuggests nothing more than tumor-specific nanoparticle 

degradation and release of drug. 35 


220 

12.4.2.1 Integrins as Targets for Imaging 

In an effort to image angiogenesis, Mulder and collaborators have created MR-detectable and 

fluorescent liposomes that are targeted to avb3 integrin on neovasculature. 36 The MR contrast 

agent Gd-DTPA-bis(stearylamide) was incorporated into PEGylated liposomes covalently 

coupled to RGD peptide. It was observed that the liposomes effectively brought about an increase 

in signal on T1-weighted images. Injection of these lipidic nanoparticles in nude mice with subcutaneously 

implanted LS174 human colon adenocarcinoma led to the ability to specifically image the 

vascular endothelial cells in the tumor. Ex vivo fluorescence confirmed that RGD liposomes 

specifically interacted with tumor endothelium associated with neovasculature. 

In a study reported by Lanza, paramagnetic molecular imaging of angiogenesis was accomplished 

in vivo with avb3-targeted lipid encapsulated perfluorocarbon nanoparticles of about 

250 nm. 37 A Vx-2 carcinoma tumor in New Zealand rabbits was imaged with a 1.5-T MRI 

system with either nontargeted or avb3-targeted nanoparticles. An eight-fold greater contrast 

enhancement was achieved with the targeted nanoparticles. In the second part of the study, paclitaxel 

loaded nanoparticles were targeted to tissue factor (TF) proteins on vascular smooth muscle 

cells with a specific TF antibody. The TF-targeted paclitaxel particles inhibited cell proliferation 

while delivery of nanoparticles to targeted cells was confirmed with fluorine spectroscopy. 

12.5 TARGETING USING FOLATE RECEPTORS 


During the past few decades, there has been great interest in the utilization of folate receptors for the 

targeted delivery of therapeutic and imaging agents. A number of delivery systems have been 

utilized for this purpose, including drug conjugates, 38–40 liposomes, 41 micelles, 42 viral vectors, 43 

and nanoparticles. 

Folate (vitamin B-9) is essential for the synthesis of nucleotides and amino acids. Two main 

groups of molecules are responsible for transport of folate molecules in vivo. Most cells in the body 

express a folate anion transporter with micro-molar affinities for folates that participates in the 

transport of coenzyme 5-methyltetrahydrofolate, the physiologic circulating reduced form of folate. 

Folate receptors (FR), by contrast, are members of the glycosylphosphatidylinositol (GPI)-linked 

membrane glycoprotein family and have high affinity for folic acid, an oxidized form of folate, and 

5-methyltetrahydrofolate, with binding affinities being in the nanomolar range (K D !1!10 K9 M) 

for the a isoform of FR. 44–46 It has been observed with few exceptions that only cells involved in 

pathologic conditions, including cancer cells, express the high affinity folate receptors. These 

receptors are able to transport folic acid, folate-bound molecules, and even particles through 

receptor-mediated endocytosis. 47,48 

FR are known to be overexpressed in various epithelial cancer cells, such as those of ovarian, 

mammary gland, colon, lung, prostate, and brain epithelial cancers, and in leukemic cells. 49–61 

Folate receptor overexpression has been correlated to poor prognosis. In addition, metastasized 

cancer cells have been found to overexpress the folate receptor to a larger degree than localized 

tumor cells. 62 This finding is of great importance. The only nonpathological tissues where FR is 

expressed are choroid plexus, placenta, lungs, thyroid, and kidney. 46,63 FR expression is limited to 

the apical (luminal) side of polarized epithelial cells, except for the cells of the proximal tubules in 

the kidney. As a consequence, FR is practically inaccessible to blood-borne folate-linked 

systems. 44,45 These characteristics make folate receptors very advantageous for targeted delivery 

of nanoparticles with high payloads of therapeutic agents, imaging agents, and even genes for 

the treatment, detection, and monitoring of cancer. What is more, the macromolecular size of 

nanoparticles will prevent gromerular filtration and the consequent exposure of kidney tissue to 

folate-targeted nanoparticles. 



12.5.1 ANTIBODIES AND FOLATE RECEPTORS 

The most common targeting moieties utilized for FR targeting include monoclonal antibodies to FR 

and folic acid. 46 To date, monoclonal antibodies to FR have not been utilized for targeted delivery 

of nanoparticles although successful experiences have been reported for radiopharmaceuticals. 64 

The benefits of targeting with folic acid are many: it is small in size, has high stability, lacks 

immunogenicity, and costs very little. 62 Folic acid conjugation at its g-carboxyl group is necessary 

for maintaining binding affinity to FR. The structure of folic acid is shown in Figure 12.1. Cellular 

uptake of drug carriers bound to folic acid is believed to be mediated by receptor mediated 

endocytosis although this process is not currently completely understood. A number of hypotheses 

have been proposed for this process, including clathrin and caveolar pathways. 63 It has been shown 

that folate-bound molecules are able to escape endosomes after receptor-mediated endocytosis 

because the process of endosomal acidification results in a conformational change in the receptor 

that facilitates folate ligand release. Consequently, folate-bound molecules provides a great opportunity 

for the delivery of pH-sensitive biopharmaceuticals. 45,62 

In most drug delivery systems investigated thus far, folic acid is incorporated to the drug 

delivery system through conjugation to a poly(ethylene glycol) (PEG) spacer utilizing wellknown 

dicyclohexylcarbodiimide/N-hydroxysuccinimide (DCC/NHS) mediated chemistry. Such 

design aims to minimize steric hindrance for optimal folate recognition. This conjugation technique, 

however, can activate both the g- and the a-carboxylic acids of folic acid. It should be said, 

though, that the g group is more reactive and is responsible for most of the linkages. 

Numerous attempts at nanoparticle targeting to the folate receptor have been reported. Many 

groups have formulated nanospheres of amphiphilic block copolymers including PEG. Park and 

collaborators reported in vitro results of the preparation and evaluation of methoxy poly(ethylene 

glycol) (PEG)-poly(3-caprolactone) (PCL) block copolymer nanospheres loaded with paclitaxel in 

which folic acid was conjugated to a modified amino-terminated PCL with a carbodiimidemediated 

reaction. 65 Dialysis was used to create these nanospheres that ranged in diameter from 

50 to 120 nm, depending on the ratio of the block copolymers. Paclitaxel loading efficiencies of up 

to 55% were reported with this system, thus significantly increasing the effective solubility of this 

agent in aqueous systems like the body. Because the folate moiety was conjugated to the hydrophobic 

end of the block copolymer, it is expected that upon nanosphere formation it will 

be localized in the inner core of the particles. However, XPS characterization demonstrated the 

presence of nitrogen-containing molecules at the surface which could only be attributed to the 

folate linker. Although the targeting effectiveness of these particles was not determined, cytotoxicity 

studies revealed that encapsulation of paclitaxel into the nanospheres reduced its 

O 

OH 

HO 

O 

O 

NH 

FIGURE 12.1 Chemical structure of folic acid. Conjugation to folic acid for folate targeting is commonly 

done at the g-carboxyl group. 


NH 

N 

HO 

N 

N 

N 

NH 2

222 

cytotoxicity. Consequently, encapsulation offered a safe alternative to direct administration of this 

chemotherapeutic agent. Further studies should evaluate whether conjugation of folic acid to 

hydrophobic end of the block copolymer results in sufficient surface availability of this 

targeting agent. 

Another approach to folate targeting of nanoparticles has been the modification of surface 

properties of polymeric nanoparticles with polymer conjugates including the targeting agent. For 

example, Kim reported on the modification of anionic poly(lactic-co-glycolic acid) (PLGA) nanoparticles 

prepared by emulsification and solvent evaporation with a cationic poly(L-lysine)poly(ethylene 

glycol)-folate (PLL–PEG–FOL) copolymer. 66 In this design, the polycation PLL 

block attaches to the PLGA nanoparticle through ionic interactions. Surface coating is achieved 

by simple incubation in an aqueous solution containing the PLL–PEG–FOL copolymer. The PEGfolate 

end of the copolymer is oriented toward the outer aqueous phase for better interaction 

between folate and the targeted cell membrane receptor. XPS characterization demonstrated the 

presence of nitrogen from folic acid on the surface of the coated nanoparticles. In vitro cellular 

studies with FITC-labeled nanoparticles revealed an increase in uptake of coated nanoparticles with 

increased conjugate-to-nanoparticle ratio in KB cells. Since a decrease in uptake was seen upon 

addition of free folic acid in the medium, the transport of nanoparticles into the cells was attributed 

to endocytosis mediated by the folate receptor. 

Dendritic polymer systems of polyamidoamine (PAMAM) with folic acid as the targeting agents 

and drug and imaging agent (methotrexate or tritium and fluorescein or 6-carboxytetramethylrhodamine) 

were prepared and tested in mouse models. 67 It was found that targeted systems, as opposed 

to nontargeted systems, slowed the rate of tumor growth and even showed a complete cure in one 

mouse. This study was conducted with twice-weekly tail vein injections. The biodistribution studies 

showed that, for a single targeted injection of nanoparticles, a very high amount of the nanoparticles 

accumulated within the tumor by day 1; this level remained high through at least day 4. 

Nanometric particles prepared from drug-polymer conjugates have also been reported. In one 

notable study, Yoo and collaborators reported on the formulation of doxorubicin-PEG-folate nanoaggregates 

encapsulating additional doxorubicin in their core. 11 These aggregates are formed 

spontaneously when an organic phase containing the copolymer and solubilized doxorubicin is 

dispersed into an aqueous phase containing triethylamine. The basic aqueous environment results in 

deprotonation of doxorubicin, causing it to form aggregates with the hydrophobic copolymer. The 

average aggregate diameter was approximately 200 nm. In vitro cellular studies showed increased 

uptake of the nano aggregates in cells expressing the folate receptor when folic acid was absent 

from cell media and increased cytotoxicity (anti-tumor efficacy) of the aggregates compared to the 

free drug in cells expressing the FR. In vivo studies in mouse xenografts (KB cells) showed that 

the nanoaggregates had superior therapeutic efficacy than both doxorubicin aggregates without the 

folic acid ligand and free doxorubicin in solution. 

Nanoparticles of temperature-responsive hydrogels conjugated to folic acid have also 

been studied with the purpose of delivering chemotherapeutic agents. Nayak reported on poly 

(N-isopropylacrylamide) (pNIPAM) nanoparticles that exhibit lower critical solution temperature 

(LCST) behavior. 68 Fluorescent agents were included in the core of these nanoparticles to facilitate 

with tracking; amine comonomers were incorporated into the other pNIPAM shell for conjugation 

to folic acid. These nanoparticles swell when their temperature falls under their LCST. Indeed, the 

nanoparticle size increased from w50 nm at 378Ctow135 nm at 258C. In vitro cell uptake studies 

showed a 10-fold increase in the intake of nanoparticles conjugated to folic acid compared to those 

without the targeting agent in KB cells (FRC). 

12.5.2 FOLATE-TARGETED NANOPARTICLES FOR GENE DELIVERY 


The use of folate-targeted nanoparticles for gene delivery has also been studied. In gene delivery, 

tissue targeting is very important for the efficacy and safety of treatment. To date, the transfection 



efficiency offered by synthetic gene delivery systems is very low compared to that of viral vectors. 

This is also true for nanoparticle-based systems. Extensive research is being done in the area to 

improve transfection efficiencies by designing better delivery systems with improved targeting 

ability, DNA stabilization, and cellular interaction. 

Mansouri reported on the formulation of nanoparticles through the complex coacervation of 

chitosan-folic acid conjugates with DNA. 69 Chitosan is a biocompatible polycation that joins to 

form a complex with DNA through electrostatic interactions and provides protection against 

nuclease degradation. Complex coacervation is an optimal preparation technique for the encapsulation 

of DNA because it avoids the use of organic solvents and high-energy ultrasonication. 

Results revealed that the integrity of plasmid DNA in the particles was maintained and that 

conjugation of chitosan to folic acid did not interfere with the electrostatic interaction between 

chitosan and DNA. As expected, the ratio of chitosan to DNA, and consequent charge ratio, had an 

effect on particle size and zeta potential. Nanoparticles smaller than 200 nm were obtained with a 

chitosan amino group to DNA phosphate group ratio of more than 2. The use of these nanoparticles 

consequently offers a promising alternative for nonviral gene therapy for the treatment of cancer 

and other diseases in which folate receptors are overexpressed. 

Nanoparticles with a poly(L-lactic acid) (PLL) core and a polyethyleneimide (PEI) surface 

conjugated to folate were utilized for delivery of plasmid DNA. 70 Folic acid was conjugated to 

the N-terminal amino group of PEI. PEI is a polycation that has been used in the past for DNA 

condensation and delivery because it protects DNA from degradation through an endosomal escape 

mechanism. Here nanoparticles were prepared through the self-assembly of the amphiphilic folate- 

PEI–PLL copolymer with DNA in an aqueous medium. Nanoparticles of approximately 100– 

150 nm in diameter and spherical shape were produced. In vitro luciferase transfection studies 

revealed that this system actually resulted in lower luciferase expression than 

PEI–DNA complexes. 

In a separate report, folate-polyethyleneglycol–distearoylphophatidylethanolamine conjugate 

(f-PEG–DSPE), 3([N-(N 0 ,N 0 -dimethylaminoethane)-carbamoyl] cholesterol, and Tween 80 were 

used to complex with DNA into cationic nanoparticles of 100–200 nm in diameter with a modified 

ethanol injection method. 71 The formulation was carried out by dissolving the lipids in ethanol and 

then removing the solvent through evaporation in the presence of water. The folate moiety, which is 

conjugated to the PEG end of one of the lipid conjugates, naturally localizes at the surface of the 

nanoparticles because of PEG migration toward the water phase. Tween 80, a nonionic surfactant, 

and PEG were incorporated with the purpose of improving the in vivo stability of the cationic 

nanoparticles through steric hindrance. The size of the nanoparticles with higher PEG content was 

maintained in the presence of serum. This suggests that these nanoparticles are better able to 

maintain their structural integrity in the presence of anionic competitors present in blood. Folate 

targeting enhanced association and transfection efficacy of nanoparticles complexed with a luciferase-encoding 

plasmid on FR(C) KB cells. The association and efficacy were reduced when folic 

acid was present in the medium, thus revealing the involvement of the folate receptor in the 

transport of the plasmid DNA into the cells. 

A possible limitation of folate targeting is the noted variability of FR expression levels not only 

between patients, but also within a single tumor. 44 In addition, it has been reported that expression 

of FR in cancerous cell lines is not representative of those one sees in vivo. 44 Consequently, 

screening protocols for FR expression will need to be utilized clinically in order to determine if 

folate-targeted therapies are appropriate. 

12.6 APPROACHES FOR CANCER TARGETING TO SPECIFIC CANCER TYPES 

In a recent review, Kim and Nie describe passive and active targeting methods and then go into 

detail on a number of active targeting techniques. 72 The active target combinations they mention 


224 

are lectin–carbohydrate, ligand–receptor and antibody–antigen. One limitation of using these 

targeting strategies is that the lectin–carbohydrate targeting systems are usually targeted to 

whole organs, making them inappropriate for targeting a cancerous part of a particular organ or 

tissue. New antibody systems show a great deal of promise. Unfortunately, they also have potentially 

harmful side effects such as advanced gastric adenocarcinoma. The latter arises when one 

attempts to target breast cancer on account of the fact that antigen-positive normal cells to the 

antibody BR96 in gastric mucosa, small intestine, and pancreas. Some of the aspects of angiogenic 

targeting mentioned here have already been covered elsewhere in this chapter. Specific targeting 

systems that are in use as cancer therapeutics are shown in Table 12.1. 8,72 

12.6.1 PROSTATE CANCER 

Recently, aptamers have been used to target nanoparticulate systems to prostate-specific membrane 

antigen, a known prostate cancer tumor marker. A model drug, rhodamine-labeled dextran, was 

encapsulated in PEGylated poly(lactic acid) nanoparticles, which were subsequently surface 

modified with a prostate specific RNA aptamer (A10). Binding of the aptamer nanoparticles to 

LNCaP cells expressing prostate specific membrane antigen in vitro was significantly enhanced 

when compared to a control of nontargeted particles. Additionally, very low binding was seen on 

nonprostate specific membrane antigen expressing cells (PC3). The nanoparticles were shown to 

both target and be taken up by the prostate cancer epithelial cells. This evidence points to the 

conclusion that this novel aptamer-based, targeted nanoparticle delivery approach can be effective. 73 

Gao and collaborators report the use of quantum dots (QD) for in vivo targeting of prostate cancer 

and imaging of tumors. 7 The core-shell CdSe–ZnS quantum dots contain tri-n-octylphosphine 

oxide (TOPO) that binds to a covering of high molecular weight ABC triblock copolymer of 

polybutylacrylate, polyethylacrylate, polymethacrylic acid, and an 8-carbon alkyl side chain. 

The complex is functionalized with PEG molecules and monoclonal antibodies to prostate-specific 

membrane antigen. Specific binding was shown for prostate cancer lines whereas low binding was 

seen to normal cells. The QD-antibody formulations were studied in vivo in a mouse model human 

prostate cancer. The nanoparticles were shown to target to the tumor both by passive and active 

antibody targeting. Sensitive and multicolor fluorescence imaging of cancer cells in vivo was 

TABLE 12.1 

Targeting Systems Utilizing Antibodies Currently in Use to Treat Cancer 

Mechanism 

Antibody 

Target Trade Name 

Agonist activity CD40, CD137 Various 

Antagonist activity CTLA4 MDX-010 

Angiogenesis inhibition VEGF Avastine 

Antibody-dependent cell-mediated cytotoxicity CD20 Rituxan w , HuMax-CD20, Zevalin w 

Inhibition of binding of extracellular growth signals HER-2/neu Herceptin 

Receptor blockage EGF receptor HuMax-EGFr 

Toxin-mediated killing CD33 Mytotarg w 

Disruption signaling HER-2/neu Pertuzumab (2C4) 

Complement-dependent cytotoxicity CD20 Rituxan w , HuMax-CD20 

Blockage ligand binding EGF receptor Erbutixe 

Antibody-dependent lysis of leukemic cells following 

cell binding 

CD52 Campath w 

Inhibits phosphorilation of tyrosine kinases EGF receptor Iressa 




accomplished in this study. Nonspecific uptake was seen in the liver and spleen with little or no 

uptake in other organs. 

12.7 TARGETED NANOPARTICLES AND IMAGING OF CANCER 

Hofmann and collaborators have evaluated the ability of super paramagnetic iron oxide nanoparticles 

(SPION) to interact with human melanoma cells in such a way that these particles could be 

selectively targeted to tumor cells ands then imaged using MRI. 74 They varied the coating placed on 

these particles and found that when comparing particles coated with poly(vinyl alcohol) (PVA), a 

vinyl alcohol/vinyl amine copolymer (amine-SPION), PVA with randomly distributed carboxylic 

groups or PVA with randomly distributed thiol groups, human cells in culture would interact strongly 

only with the amine-SPION particles. Furthermore, these particles showed the lowest cytotoxicity. 

This human study involved intravenous administration of Combidex (Advanced Magnetics, 

Inc., ferumoxtran-10, a molecular imaging agent of iron oxide nanoparticles and a dense packing of 

dextran derivatives) to 18 men ages 21–46 with diagnosed testicular cancer. 75 The Combidex was 

imaged using MRI. From this study, it seems evident that those lymph nodes with a higher signal 

were classified as being malignant. However, based on the information from Advanced Magnetics, 

these particles should accumulate selectively in noncancerous lymph node tissue. The particles are 

still experimental and rightly so. 

Although treatment of cancer with targeted nanoparticles is an important goal, more accurate 

imaging of cancer is needed to allow for the optimal treatment for each patient. Towards that end, 

a considerable amount of research is underway with various imaging techniques to establish more 

accurate determination of the presence and extent of cancer growth and metastases. Because magnetic 

resonance imaging (MRI) is a widely used imaging technique, much work is currently being done to 

develop targeted imaging agents for MRI. Some of these involve paramagnetic and superparamagnetic 

iron oxides due to their ability to affect water relaxation times T1 and T2. Gasco and collaborators 

have prepared solid lipid nanoparticles containing Endorem, superparamagnetic iron oxide nanoparticles 

(Guebert and Advanced Magnetics), using either a multiple emulsion technique or an oil in 

water emulsion technique. 76 Although the loading rates achieved were less than 1 wt% iron, it was 

possible to detect and image in vivo in rats. Incorporation of the Endorem in the SLN allowed passage 

across the blood brain barrier, passage which was not possible with Endorem alone. 

Often development of nanoparticle systems is a two-pronged approach involving both drug and 

imaging agent. If a targeting system is successful, one should be able to enhance the imaging of a 

cancer and then kill it with the same system. One such study involved the preparation of glycol– 

chitosan nanoaggregates to which either fluorescein isothiocyanate (FITC) or doxorubicin (Dox) 

was conjugated. 12 Based on a single tail-vein injection of FITC-conjugates, levels remained high for 

eight days and gradually increased in the tumors of rats with II45 mesothelioma cells, in the kidney, 

and to a lesser extent in the spleen. Meanwhile, levels decreased in other organs. The liver showed 

some accumulation at day 3 but was significantly lower at day 8 relative to days 1 and 3. The 

performance of these systems, as evidenced by a decrease in tumor volume, was excellent with a 

consistent decrease in tumor volume after day 13 when a tail-vein injection of Dox-nanoaggregates 

is given at days 13 and 19. 

12.8 OTHER TARGETS FOR CANCER 

A new approach to targeting is the use of lectins, which are plant proteins that specifically recognize 

cell surface carbohydrates. The latter function as selective cancer-cell-targeting agents. PLGA 

nanoparticles of mean diameter 331 nm incorporating isopropyl myristate were used to deliver 

paclitaxel to malignant A549 and H1299 and normal CCL-186 pulmonary cells in vitro by means of 

wheat germ agglutinin lectin as the targeting molecule. The in vitro cytotoxicity against A549 and 


226 

H1299 cells was significantly increased with wheat germ agglutinin-conjugated PLGA nanoparticles 

containing IPM and paclitaxel compared to controls. 77 In a subsequent study, Mo and Lim 

evaluated in vivo efficacy of the same nanoparticles in a SCID mouse model injected with an A549 

tumor nodule. 78 One injection of wheat germ agglutinin-conjugated PLGA nanoparticles 

containing IPM and a paclitaxel dose of 10 mg/kg inhibited tumor growth without appreciable 

weight loss. Tumor doubling was increased to 25 days compared to 11 days for conventionally 

formulated paclitaxel. 

12.9 AVIDIN AND BIOTIN TARGETING 

Although it is by far the most widely utilized polymer for surface modification of nanoparticles, 

PEG is not the only compound that can be included at the surface of nanoparticles. Nor is it 

necessary to achieve active targeting. Saltzman has recently reported a method for incorporating 

avidin-fatty acid conjugates into the surface of PLGA nanoparticles. 79 This method resulted in 

avidin at the surface of the nanoparticles that remained active for weeks. The ability of these 

nanoparticles to target to biotin was verified by targeting of the nanoparticles to biotinylated 

agarose beads. Not only could this system be used for targeted delivery; it can also be utilized to 

selectively modify surfaces for tissue engineering. 

In another such example, Hunziker has prepared biotin-functionalized (poly(2-methyloxazoline)-b-poly(dimethylsiloxane)-b-poly(2-methyl-oxazoline) 

triblock copolymers. 80 The 

biotinylated targeting agents were added using streptavidin as a coupling agent. Uptake of these 

“nanocontainers” was seen in the presence of the target receptor, the macrophage scavenger 

receptor SRA1, but not in the absence of this target. 


The amount of research in targeted, polymeric nanoparticles for cancer imaging and therapy has 

increased dramatically in the past 5–10 years. Seeing actual products using targeted therapies has 

no doubt fueled that work. In the next decade, we will certainly see products, whether with 

polymeric nanoparticles or some other type of delivery system, using folate receptors and carrying 

imaging agents. All of these technologies, driven by the fields of fundamental immunology, 

biochemistry, polymer chemistry, and biomedical engineering, are bringing us closer to the time 

when cancer may be treated on an individual basis. One patient’s diagnosis and treatment will be 

unique to her condition and will be the most effective treatment possible for her. Until other 

scientists determine how to stop cancer from occurring, those mentioned in this chapter and 

many more besides them are doing their best to eliminate cancer. 

REFERENCES 


1. Ferlay, J. et al., GLOBOCAN 2002: Cancer Incidence, Mortality and Prevalence Worldwide, IARC 

CancerBase No. 5, Version 2.0., IARC Press, Lyon, 2004. 

2. Society, A. C., Cancer Facts and Figures 2005, American Cancer Society, Atlanta, 2005. 

3. Yokoyama, M. and Okano, T., Targetable drug carriers: Present status and a future perspective, 


4. Torchilin, V. P., Strategies and means for drug targeting: An overview, In Biomedical Aspects of Drug 

Targeting, Muzykantov, V. and Torchilin, V., Eds., Kluwer Academic, Norwell, pp. 3–26, 2002. 

5. McDonald, D. M. and Baluk, P., Significance of blood vessel leakiness in cancer, Cancer Research, 

62, 5381–5385, 2002. 

6. Conner, S. D. and Schmid, S. L., Regulated portals of entry into the cell, Nature, 422, 37–44, 2003. 

7. Gao, X. et al., In vivo cancer targeting and imaging with semiconductor quantum dots, Nature 

Biotechnology, 22(8), 969–976, 2004. 





9. Luo, Y. et al., Targeted delivery of doxorubicin by HPMA copolymer-hyaluronan bioconjugates, 

Pharmaceutical Research, 19, 396–402, 2002. 

10. Mastrobattista, E. et al., Targeted liposomes for delivery of protein-based drugs into the cytoplasm of 

tumor cells, Journal of Liposome Research, 12, 57–65, 2002. 

11. Yoo, H. S. and Park, T. G., Folate receptor targeted biodegradable polymeric doxorubicin micelles, 


12. Vauthier, C. et al., Drug delivery to resistant tumors: The potential of poly(alkly cyanoacrylate) 

nanoparticles, Journal of Controlled Release, 93, 151–160, 2003. 

13. Peracchia, M. T., Stealth nanoparticles for intravenous administration, S.T.P. Pharma, 13(3), 

155–161, 2003. 

14. Shenoy, D. B. and Amiji, M. M., Poly(ethylene oxide)-modified poly(e-caprolactone) nanoparticles 

for targeted delivery of tamoxifen in breast cancer, International Jounal of Pharmaceutics, 293, 

261–270, 2005. 

15. Avgoutsakis, K. et al., PLGA–mPEG nanoparticles of cisplatin: In vitro nanoparticles degradation, 

in vitro drug release and in vivo drug residence in blood properties, Journal of Controlled Release, 79, 

125–135, 2002. 

16. Machida, Y. et al., Efficacy of nanoparticles containing irinotecan prepared using poly(DL-lactic acid) 

and poly(ethylene glycol)-poly(propylene glycol)-poly(ethylene glycol) against M5076 tomur in the 

early liver metastatic stage, S.T.P. Pharma, 13(4), 225–230, 2003. 

17. Brigger, I. et al., Poly(ethylene glycol)-coated hexadecylcyanoacrylate nanospheres display a 

combined effect for brain tumor targeting, Journal of Pharmacology and Experimental Therapeutics, 

303, 928–936, 2002. 

18. Reddy, L. H., Sharma, R. K., and Murthy, R. S. R., Enhanced tumour uptake of doxorubicin loaded 

poly(butyl cyanoacrylate) nanoparticles in mice bearing Dalton’s lymphoma tumour, Journal of Drug 

Targeting, 12(7), 443–451, 2004. 

19. Kaul, G. and Amiji, M., Biodistribution and targeting potential of poly(ethylene glycol)-modified 

gelatin naoparticlea in subcutaneous murine tumor model, Journal of Drug Targeting, 12(9–10), 

585–591, 2004. 

20. Folkman, J., Angiogenic-dependent diseases, Seminars in Oncology, 28(6), 536–542, 2001. 

21. Lyden, D. et al., Impaired recruitment of bone-marrow-derived endothelial and hematopoietic 

presursor cells blocks tumor angiogenesis and growth, Nature Medicine, 7(11), 1194–1201, 2001. 

22. Folkman, J., Vascularization of tumors, Scientific American, 234(5), 59–76, 1976. 

23. Folkman, J., Fundamental concepts of the angiogenic process, Current Molecular Medicine, 3, 

643–651, 2003. 

24. Hanahan, D. and Folkman, J., Patterns and emerging mechanisms of the angiogenic switch during 

tumorgenesis, Cell, 86, 353–364, 1996. 

25. Carmeliet, P. and Jain, R. K., Angiogenesis in cancer and other diseases, Nature, 407, 249–257, 2000. 

26. Leenders, W. P. J., Targetting VEGF in anti-angiogenic and anti-tumour therapy: Where are we now?, 

International Journal of Experimental Pathology, 79, 339–346, 1998. 

27. Claffey, K. P. et al., Expression of vascular permeasbility factor/vascular endothelial growth factor by 

melanoma cells increases tumor growth, angiogenesis, and experimental metastasis, Cancer Research, 

56, 172–181, 1996. 

28. Potgens, A. J. G. et al., Analysis of the tumor vasculature and metastatic behavior of xenografts of 

human melanoma cells lines transfected with vascular permeability factor, American Journal of 

Pathology, 148(4), 1203–1217, 1996. 

29. Li, L. et al., A novel antiangiogenesis therapy using an integrin antagonist or anti-FLK-1 antibody 

coated with 90 Y-labeled nanoparticles, International Journal of Radiation Oncology Biology Physics, 

58(4), 1215–1227, 2004. 

30. Brooks, P. C., Clark, R. A. F., and Cheresh, D. A., Requirement of vascular integrin a vb 3 for 

angiogenesis, Science, 264, 569–571, 1994. 

31. Varner, J. A. and Cheresh, D. A., Integrins and cancer, Current Opinion in Cell Biology, 8, 724–730, 

1996. 


228 



2404–2407, 2002. 

33. Schiffelers, R. M. et al., Cancer siRNA therapy by tumor selective delivery with ligand-targeted 

sterically stabilized nanoparticle, Nucleic Acids Research, 32(19), e149, 2004. 

34. Kopelman, R. et al., Multifunctional nanoparticle platforms for in vivo MRI enhancement and photodynamic 

therapy of a rat brain cancer, Journal of Magnetism and Magnetic Materials, 293, 404–410, 

2005. 

35. Bibby, B. C. et al., Pharmacokinetics and biodistribution of RGD-targeted doxorubicin-loaded nanoparticles 

in tumor-bearing mice, International Jounal of Pharmaceutics, 293, 281–290, 2005. 

36. Mulder, W. J. M. et al., MR molecular imaging and fluorescence microscopy for identification of 

activated tumor endothelium using a bimodal lipidic nanoparticle, The FASEB Journal, 19, 

2008–2010, 2005. 

37. Lanza, G. M., Magnetic resonance molecular imaging and targeted drug delivery with site-specific 

nanoparticles. In BioMEMS and Biomedical Nanotech World, Columbus, OH, 2002. 

38. Leamon, C. P. et al., Synthesis and biological evaluation of EC72: A new folate-targeted chemotherapeutic, 

Bioconjugate Chemistry, 16, 803–811, 2005. 

39. Lee, J. W. et al., Synthesis and evaluation of taxol-folic acid conjugates as targeted antineoplastics, 

Bioorganic Medicinal Chemistry, 10, 2397–2414, 2002. 

40. Paranjpe, P. V. et al., Tumor-targeted bioconjugate based delivery of camptothecin: Design, sythesis 

and in vitro evaluation, Journal of Controlled Release, 100, 275–292, 2004. 

41. Gabizon, A. A., Liposome circulation time and tumor targeting: Implications for cancer 

chemotherapy, Advanced Drug Delivery Reviews, 16, 285–294, 1995. 

42. Lee, D. H., Kim, D. I., and Bae, Y. H., Doxorubicin loaded pH-sensitive micelle targeting acidic 

extracellular pH of human ovarian A2780 tumor in mice, Journal of Drug Targeting., 13(7), 391–397, 

2005. 

43. Douglas, J. T. et al., Targeted gene delivery by tropism-modified adenoviral vectors, Nature Biotechnology, 

14, 1574–1578, 1996. 

44. Elnakat, H. and Ratnam, M., Distribution, functionality and gene regulatio of folate receptor isoforms: 

Implications in targeted therapy, Advanced Drug Delivery Reviews, 56, 1067–1084, 2004. 

45. Leamon, C. P. and Reddy, J. A., Folate-targeted chemotherapy, Advanced Drug Delivery Reviews, 56, 

1127–1141, 2004. 

46. Sudimack, J. and Lee, R. J., Targeted drug delivery via the folate receptor, Advanced Drug Delivery 

Reviews, 41, 147–162, 2000. 

47. Antony, A., The biological chemistry of folate receptors, Blood, 79, 2807–2820, 1992. 

48. Kamen, B. and Capdevila, A., Receptor-mediated folate accumulation is regulated by the cellular 

folate content, Proceedings of the National Academy of Sciences of the United State of America, 88, 

5983–5987, 1986. 

49. Mattes, M. et al., Patterns of antigen distribution in human carcinomas, Cancer Research, 50, 

880s–884s, 1990. 

50. Coney, L. et al., Distribution of the folate receptor GP38 in normal and malignant cell lines and 

tissues, Cancer Research, 52, 3396–3401, 1991. 

51. Weitman, S. et al., Distribution of the folate receptor GP38 in normal and malignant cell lines and 

tissues, Cancer Research, 52, 3396–3401, 1992. 

52. Ross, J., Chaadhuri, P., and Ratnam, M., Differential regulation of folate receptor isoforms in normal 

and malignant cell lines. Physiologic and clinical implications, Cancer, 73, 2432–2443, 1994. 

53. Weitman, S. et al., Cellular localization of thr folate receptor: Potential role in drig toxicity and folate 

homeostasis, Cancer Research, 52, 6708–6711, 1992. 

54. Weitman, S., Frazier, K., and Kamen, B., The folate receptor in central nervous system malignancies 

of childhood, Journal of Neurooncology, 21, 107–112, 1994. 

55. Garin-Chesa, P. et al., Trophoblast and ovarian cancer antigen LK26. Sensitivity and specificity in 

immunopathology and molecular identification as a folate-binding protein, American Journal of 

Pathology, 142, 557–567, 1993. 

56. Toffoli, G. et al., Overexpression of folate binding protein in ovaian cancers, International Journal of 

Cancer, 74, 193–198, 1997. 



57. Holm, J. et al., High-affinity folate binding in human choroid plexus. Characterization of radioligand 

binding, immunoreactivity, molecular heterogeneity and hydrophobic domain of the binding protein, 

Biochemical Journal, 280, 267–271, 1991. 

58. Shen, F. et al., Identification of a novel folate receptor, a truncated receptor and receptor type B in 

hematopoietic cells: cDNA cloning, expression, immunoreactivity, and tissue specificity, Biochemistry, 

33, 1209–1215, 1994. 

59. Sadasivan, E. et al., Characterization of multiple forms of folate-binding protein from himan leukemia 

cells, Biochimica et Biophysica Acta, 882, 311–321, 1986. 

60. Sadasivan, E. et al., Purification, properties, and immunological characterization of folate-binding 

proteins from human leukemia cells, Biochimica et Biophysica Acta, 925, 36–47, 1987. 

61. Morikawa, J. et al., Identification of signature genes by microarray for acute myeloid leukemia 

without maturation and acute promyelocytic leukemia with t(15;17)(q22;q12)(PML/RARalpha), 

International Journal of Oncology, 23, 617–625, 2003. 

62. Hilgenbrink, A. R. and Low, P. S., Folate receptor-mediated drug targeting: From therapeutics to 

diagnostics, Journal of Pharmaceutical Sciences, 94(10), 2135–2146, 2005. 

63. Sabharanjak, S. and Mayor, S., Folate receptor endocytosis and trafficking, Advanced Drug Delivery 

Reviews, 56, 1099–1109, 2004. 

64. Coliva, A. et al., 90Y Labeling of monoclonal antibody MOv18 and preclinical validation for radioimmunotherapy 

of human ovarian carcinomas, Cancer Immunology, Immunotherapy, 54, 1200–1213, 

2005. 

65. Park, E. K., Lee, S. B., and Lee, Y. M., Preparation and characterization of methoxy poly(ethylene 

glycol)/poly(3-caprolactone) amphiphilic block copolymeric nanospheres for tumor-specific folatemediated 

targeting of anticancer drugs, Biomaterials, 26, 1053–1061, 2005. 

66. Kim, S. H. et al., Target-specific cellular uptake of PLGA nanoparticles coated with poly(L-lysine)poly(ethylene 

glycol)-folate conjugate, Langmuir, 21, 8852–8857, 2005. 

67. Kukowska-Latallo, J. F. et al., Nanoparticle targeting of anticancer drug improves therapeutic 

response in animal model of human epithelial cancer, Cancer Research, 65(12), 5317–5324, 2005. 

68. Nayak, S. et al., Folate-mediated cell targeting and cytotoxicity using thermoresponsive microgels, 

Journal of the American Chemistry Society, 126, 10258–10259, 2004. 

69. Mansouri, S. et al., Characterization of folate–chitosan–DNA nanoparticles for gene therapy, Biomaterials, 

27(9), 2060–2065, 2006. 

70. Wang, C.-H. and Hsiue, G.-H., Polymer-DNA hybrid nanoparticles based on folate-polyethyleneimine-block-poly(L-lactide), 


71. Hattori, Y. and Maitani, Y., Enhanced in vitro DNA transfection efficiency by novel folate-linked 

nanoparticles in human prostate cancer and oral cancer, Journal of Controlled Release, 97, 173–183, 

2004. 

72. Kim, G. J. and Nie, S., Targeted cancer nanotherapy, Nanotoday, August, 28–33, 2005. 

73. Farokhzad, O. C. et al., Nanoparticles-aptamer bioconjugates: A new approach for targeting prostate 

cancer cells, Cancer Research, 64, 7668–7672, 2004. 

74. Petri-Fink, A. et al., Development of functionalized superparamagnetic iron oxide nanoparticles for 

interaction with human cancel cells, Biomaterials, 26, 2685–2694, 2005. 

75. Harisinghani, M. G. et al., A pilot study of lymphotrophic nanoparticle-enhanced magnetic resonance 

imaging technique in early stage testicular cancer: A new method for noninvasive lymph node 

evaluation, Urology, 66, 1066–1071, 2005. 

76. Peira, E. et al., In vitro and in vivo study of solid lipid nanoparticles loaded with superparamagnetic 

iron oxide, Journal of Drug Targeting, 11(1), 19–24, 2003. 

77. Mo, Y. and Lim, L.-Y., Preparation and in vitro anticancer activity of wheat germ agglutinin (WGA)conjugated 

PLGA nanoparticles loaded with paclitaxel and isopropyl myristate, Journal of Controlled 

Release, 107, 30–42, 2005. 

78. Mo, Y. and Lim, L., Paclitaxel-loaded PLGA nanoparticles: Potentiation of anticancer activity by 

surface conjugation with wheat germ agglutinin, Journal of Controlled Release, 108, 244–262, 2005. 

79. Fahmy, T. M. et al., Surface modification of biodegradable polyesters with fatty acid conjugates for 

improved drug targeting, Biomaterials, 26, 5727–5736, 2005. 

80. Broz, P. et al., Cell targeting by a generic receptor-targeted polymer nanocontainer platform, Journal 



13 

CONTENTS 

Long-Circulating Polymeric 

Nanoparticles for Drug and 

Gene Delivery to Tumors 

Sushma Kommareddy, Dinesh B. Shenoy, 

and Mansoor M. Amiji 

13.1 Introduction ....................................................................................................................... 231 

13.2 Polymeric Nanoparticles ................................................................................................... 232 

13.3 Nanoparticles for Drug Delivery ...................................................................................... 233 

13.4 Nanoparticles for Gene Delivery ...................................................................................... 233 

13.5 Long-Circulating Nanoparticles ........................................................................................ 234 

13.6 Illustrative Examples of Polymeric Nanoparticles for Drug Delivery............................. 235 

13.6.1 Polyethylene Oxide-Modified Poly(b-Amino Ester) Nanoparticles .................. 235 

13.6.2 Poly(Ethylene Oxide)-Modified Poly(3-Caprolactone) Nanoparticles............... 236 

13.7 Illustrative Examples of Polymeric Nanoparticles for Gene Delivery ............................ 237 

13.7.1 Polyethylene Glycol-Modified Gelatin Nanoparticles........................................ 237 

13.7.2 Polyethylene Glycol-Modified Thiolated-Gelatin Nanoparticles ....................... 238 

13.8 Concluding Remarks ......................................................................................................... 238 


References .................................................................................................................................... 239 


During the past few decades, there have been extensive efforts in the treatment and cure of cancer 

that is still the second leading cause of death next to the cardiovascular diseases. With advances in 

research techniques, there has been a surge of therapeutic agents in the form of proteins and nucleic 

acids as a source of new chemical entities for the treatment of cancer. 1–7 However, effective 

delivery of these novel agents to the target tissue has always been a problem. 

The therapeutic agents used in cancer treatment are generally administered in the systemic 

circulation. The drug carrier, therefore, must overcome physiological barriers to reach the tumor 

cell in sufficient concentrations and to reside for the necessary duration to exert the pharmacological 

effect. These barriers include transport of the drugs within the blood vasculature and 

transport from the vasculature into the surrounding tumor tissues and through the interstitial spaces 

within the tumor. Solid tumors are characterized by vasculature that is heterogeneous in size and 

distribution, having a central avascular/necrotic region and vascularized peripheral region with 


231

232 

discontinuous endothelium in the microvessels. Depending on the anatomic region of the tumor, the 

pore size of the endothelial junctions is found to vary from 100 to 780 nm with a mean of 

approximately 400 nm. 8–11 Tumor vasculature is also characterized by a lack of lymphatic drainage. 

12–14 In addition to the complexity of tumor physiology, the development of multi-drug 

resistance (MDR) in tumor cells exacerbates the problem of achieving selective toxicity in 

tumor tissues. 

Typically, a tumor consists of neoplastic cells, stromal cells, and the extracellular matrix and is 

associated with blood vessels, nerves, and immune cells that form an integral part of the stroma. 15,16 

The tumor also contains leukocytes, fibroblasts, and other extracellular elements. The majority of 

the infiltrated leukocytes are macrophages, more popularly referred to as tumor-associated macrophages 

(TAM). They have increased phagocytic activity and play an active role in tumor 

progression. The presence of these activated macrophages in large numbers might lead to enhanced 

clearance of a drug-carrying vector from the tumor microenvironment. 15–18 

13.2 POLYMERIC NANOPARTICLES 


Sub-micronic particles prepared from polymers are becoming increasingly popular for the delivery 

of drugs or genes to tumor tissues. When compared to all other colloidal delivery systems, nanoparticles 

have better stability in plasma and higher encapsulation efficiency, and they are amenable 

to large-scale preparation. These nanoparticles can be used to increase the solubility of hydrophobic 

drugs, lower the toxicity of drugs with a high therapeutic index, and enhance the stability of the 

payload, in the case of DNA, by protecting it from degradation by extracellular enzymes. Furthermore, 

these particles permit controlled release of the payload at the target site at relatively low 

doses. 19 In addition, there is a wide range of polymeric materials whose physicochemical properties 

can be tailored to achieve polymeric nanoparticles of the required nature. 

Passive targeting of these particles at the tumor sites is generally achieved by the enhanced 

vascular permeability and lack of lymphatic drainage, together termed the enhanced permeability 

and retention effect (EPR) (Figure 13.1). 12–14 The accumulation of polymeric nanoparticles 

carrying drugs is dependent on the physical chemistry of the polymer, including molecular 

weight, surface charge, nature of the polymer, etc. The EPR is one of the main reasons for the 

success of polymeric nanoparticles in targeting tumors. Site-specific or active targeting of these 

particles can be achieved by conjugating with targeting moieties or ligands that are specific to the 

tumor cells. These targeting moieties present on the polymer backbone are used to exploit 

the differences between tumor cells and normal cells through receptor-mediated endocytosis. 

Active targeting, in particular, can be used to overcome the obstacle of tumor metastasis and to 

increase the specificity and efficacy of the polymeric carrier system. Transferrin, folate, epidermal 

growth factor, and argenine–glycine–aspartic acid (RGD) tripeptide are some of the moieties that 

are used for targeting tumor cells. 7,20–26 

The major drawback to the use of nanoparticles in cancer therapy is their initial burst release of 

the drug upon administration into systemic circulation that is followed by slow controlled release 

of the encapsulated drug. The polymers used in the preparation of nanoparticles generally lack the 

ability to encapsulate both hydrophilic and hydrophobic drugs in the same polymer system (i.e., the 

hydrophilic drugs have to be encapsulated in a hydrophilic polymer). Some of the synthetic polymers 

have cytotoxicity issues associated with them. However, this difficulty could be overcome by 

using biosynthetic polymers that are biodegradable and modified to have physicochemical properties 

similar to those of the synthetic polymers. The use of polymeric nanoparticles for systemic 

delivery must be limited to non-cationic or surface-modified polymers to prevent adsorption of 

plasma proteins onto the surface of the nanoparticles. Furthermore, some of the naturally occurring 

polymers have immunogenic reactions when injected into blood; such reactions could be prevented 

through the use of non-immunogenic biopolymers now available. 


Long-Circulating Polymeric Nanoparticles for Drug and Gene Delivery to Tumors 233 

Extravascular Space 

PEG-Modified Polymeric Nanoparticles 

Blood flow 

Extravascular Space 

13.3 NANOPARTICLES FOR DRUG DELIVERY 

The majority of the chemotherapeutic agents used in cancer therapy are low molecular weight drugs 

with a high volume of distribution, leading to the presence of toxic compounds throughout the 

body. These drugs must be given in high doses in order to reach therapeutic levels in tumor tissues. 

In the process, healthy tissues are exposed to cytotoxic drugs, resulting in side effects such as bone 

marrow suppression, alopecia, anorexia, etc. The low molecular weight anti-cancer agents are also 

rapidly cleared from the body. Continuous treatment of tumor cells with these drugs would exacerbate 

the problem of MDR that is one of the main reasons for the failure of chemotherapy. 19 

Numerous efforts in improving cancer chemotherapy have focused on increasing the therapeutic 

index and reducing the non-selective cytotoxicity of the drugs in use. Macromolecular 

carriers such as polymer conjugates, liposomes, microspheres, and nanoparticles are used for 

this purpose. Of these, the nanoparticles have been promising, mainly because of their advantages 

over other novel drug delivery vehicles such as their safety, enhanced stability, possibility of 

tailoring and surface modification, and industrial scale-up. 27 Nanoparticles prepared from 

polymer–drug conjugates have altered pharmacokinetic distribution and increased pharmacological 

activity at tolerable doses. This is due mainly to the controlled release of the drug and reduced renal 

clearance of the low molecular weight drugs. 

The concept of targeting drugs at the site of action was first described as “magic bullets” by 

Paul Ehrlich. 28 Later, Ringsdorf presented a model of the polymer–drug conjugate that could be 

used to improve chemotherapy drugs for cancer. 29 The conjugation or encapsulation of these low 

molecular weight drugs within polymeric nanoparticles has distinct advantages: Increased plasma 

half-life of low molecular weight drugs; increased solubility of hydrophobic drugs; and controlled 

release of the drug at lower doses. The targeting of these nanoparticles to the tumor tissues is 

achieved either passively by the EPR effect or actively by surface conjugating a tumor-specific 

ligand. 30 

13.4 NANOPARTICLES FOR GENE DELIVERY 

Blood Capillary 

Endothelial Cells 

Leaky Vasculature 

FIGURE 13.1 Schematic of enhanced permeability and retention effect. Passive targeting to the tumor cells is 

achieved by the extravasation of the polymeric nanoparticles through the leaky vasculature of the blood 

capillaries in the tumor. The leaky vasculature, along with the lack of lymphatic drainage, is called the 

EPR effect and is used to passively target nanocarriers to the tumor. 

Due to the complexity and heterogeneity in cancer, gene therapy can provide a unique approach for 

treatment. Research on gene therapy focuses on the successful in vivo transfer of genes to the target 

tissue for the sustained expression of the genes of dysfunction. Various vectors, both viral and nonviral 

in origin, are used for this purpose. 31–34 The limitations associated with viral vectors, namely 

issues of integration with the host genome, self-replication, recombination potential, and 


234 

immunogenicity, have encouraged researchers to seek alternatives such as nanoparticles, liposomes, 

and other cationic complexes. 35–37 Despite the low transfection efficiency associated with 

these non-viral vectors, they have the advantages of safety and high encapsulation efficiency. The 

DNA complexes of cationic liposomes and polymers have high transfection efficiency in vitro; 

however, owing to their toxicity and rapid clearance, these formulations have limited efficiency 

in vivo. The cationic liposomes and polymers, being positively charged, form aggregates with the 

negatively charged serum proteins and opsonins, resulting in enhanced phagocytosis and clearance 

from the blood circulation. 38 

In addition to cationic polymers, several other natural and synthetic polymers have been used in 

preparing DNA-encapsulated nanoparticles. Polymeric nanoparticles are increasingly popular, 

owing to such advantages as their small size, ease of production, and administration. Several 

polymers have been investigated as vectors for gene delivery applications. There has been a fair 

amount of success in reducing immunogenicity and cytotoxicity with the concomitant enhancement 

of the efficiency of transfection with these polymers. Besides protecting the DNA from degradation, 

the nanoparticles enhance the targeting of tumor tissues. 38 

13.5 LONG-CIRCULATING NANOPARTICLES 

Long-circulating nanoparticles can be created through surface modification of conventional nanoparticles 

with water-soluble polymers such as polyethylene glycol (PEG) or polyethylene oxide 

(PEO) (Figure 13.2). The hydrophilic nature of these surface modifiers minimizes the interactions 

between the nanoparticles and plasma proteins (opsonins), resulting in reduced uptake by the 

reticulo-endothelial system (RES). The major outcome of modification with PEG or other hydrophilic 

flexible polymers is a significant increase in circulation time, the advantages of which include 

maintenance of optimal therapeutic concentration of the drug in the blood after a single administration 

of the drug carrier, increased probability of extravasation and retention of the colloidal 

carrier in areas of discontinuous endothelium, and enhancement in targetability of the system by use 

of a target-specific ligand. 39 

The protective action (Stealth w property) of PEG is mainly due to the formation of a dense, 

hydrophilic cloud of long polyethylene chains on the surface of the colloidal particle that reduces 

the hydrophobic interactions with the RES. The tethered or chemically anchored PEG chains can 

(a) 

PEG PPO 

(b) 


PEO PEO 

FIGURE 13.2 Schematic of long-circulating polymeric nanoparticles. The polymeric nanoparticles are made 

long-circulating by surface modification. The nanoparticles prepared from a hydrophilic polymer are modified 

using poly(ethylene glycol) (a), and the nanoparticles prepared from hydrophobic polymers are modified using 

Pluronic w , a triblock co-polymer of poly(ethylene oxide) and poly(propylene oxide) (b). 



undergo spatial conformations, thereby preventing the opsonization of particles by the RES of the 

liver and spleen and improving the circulation time of molecules and particles in the blood. The 

greater the flexibility of the polymer, the greater the total number of possible conformations and 

transitions from one conformation to another. 40–43 Water molecules form a structured shell through 

hydrogen bonding to the ether oxygens of PEG. The tightly bound water around PEG chains forms a 

hydrated film around the particle and prevents protein interactions. 44 Furthermore, PEGylation may 

also increase the hydrodynamic size of the particles, decreasing their clearance through the kidneys, 

renal filtration being dependent on molecular mass and volume. This would ultimately result in an 

increase in the circulation half-life of the particles. 45,46 

The size, molecular weight, and shape of the PEG fraction and the linkage used to connect it to 

the entity of interest determine the consequences of PEGylation in relation to protein adsorption 

and pharmacokinetics such as volume of distribution, circulation time, and renal clearance. When 

formulated into colloidal particles, the PEG density on the colloidal surface can be changed by 

using PEG of appropriate molecular weight (PEG chain length) and molar ratio (the grafting 

efficiency). Longer PEG chains offer greater steric influence around the colloidal entity, similar 

to increased grafting density with shorter PEG chains. Longer PEG chains may also collapse onto 

the nanoparticle surface, providing a hydrophilic shield. 40 

Besides PEG, other hydrophilic polymers, including polyvinyl alcohol, polyacryl amide, polyvinyl 

pyrrolidone, poly-[N-(2-hydroxypropyl)methacrylamide], polysorbate-80, and block 

co-polymers such as poloxomer (Pluronic w ) and poloxamine (Tetronic w ), are also being used to 

modify the physicochemical properties of the colloidal carriers. 32,47,48 

The polymeric nanoparticles modified with PEG or PEO are mainly used to passively target 

tumors through the EPR effect. The surface-modified long-circulating polymeric nanoparticles are 

used to deliver both genes and drugs to the tumor tissues. Some of the polymeric nanoparticulate 

systems that were developed by this group will be discussed in the sections that follow. 

13.6 ILLUSTRATIVE EXAMPLES OF POLYMERIC NANOPARTICLES 

FOR DRUG DELIVERY 

13.6.1 POLYETHYLENE OXIDE-MODIFIED POLY(b-AMINO ESTER) NANOPARTICLES 

A representative biodegradable, hydrophobic poly(b-amino ester) (PBAE) with pHsensitive 

solubility properties was synthesized by conjugate addition of 4,4 0 -trimethylenedipiperidine with 

1,4-butanediol diacrylate, developed in Professor Robert Langer’s lab at Massachusetts Institute of 

Technology. The paclitaxel-loaded nanoparticles (150–200 nm) prepared from PBAE were 

modified with Pluronic w F-108 (poloxamer 407), a triblock copolymer of polyethylene oxide/polypropylene 

oxide/polyethylene oxide (PEO/PPO/PEO). The PPO segment of the triblock polymer 

attaches to the hydrophobic surface of the nanoparticles, and the hydrophilic PEO segment contributes 

to the stealth properties of the polymeric nanoparticles. The pH-sensitive nature of the particles 

prepared from PBAE has already been shown by in vitro release studies carried out in the presence 

of buffers of pH ranging from 5.0 to 7.4 and was found to rapidly degrade in a medium of pH less 

than 6.5. Therefore, these nanoparticles were expected to readily release their contents within the 

acidic tumor microenvironment and in the endosomes and lysosomes of the cells upon internalization. 

This was confirmed by the in vitro cellular uptake of the PEO–PBAE nanoparticles 

encapsulated with tritiated [ 3 H]-paclitaxel by human breast adenocarcinoma cells (MDA- 

MB231). 49,50 

The biodistribution of these PEO-modified PBAE nanoparticles was carried out by encapsulating 

a lipophilic form of the radionuclide indium-111 ( 111 indium oxine). Following tail vein 

injection in nude mice bearing a human ovarian xenograft, the radiolabeled PEO-modified 

PBAE nanoparticles were found to accumulate in the highly perfused organs such as the liver, 

spleen, and lungs with greater entrapment in the microvasculature of the lungs during the initial 


236 


time points. The increasing concentrations in the kidney also indicate that the nanoparticles, once 

internalized, were disintegrated and eliminated through the kidney. The plasma half-life of the 

unmodified nanoparticles was reported to be one to ten minutes. By virtue of surface modification 

with PEO, the PBAE nanoparticles were shown to have improved circulation times, resulting in a 

mean residence time in the systemic circulation of 21 h. The paclitaxel-encapsulated PEO–PBAE 

nanocarriers were found to deliver the drug efficiently to solid tumors, resulting in a 5.2-fold and 

23-fold higher concentration of the drug at one hour and five hours post-administration relative to 

the solution form of the drug. 40 From the tumor accumulation of the paclitaxel-loaded ( 3 H-labeled) 

nanoparticles, it is evident that the pH-sensitive PEO–PBAE nanoparticle formulations can deliver 

significantly higher concentrations of the drug into the tumor than the solution form. 

13.6.2 POLY(ETHYLENE OXIDE)-MODIFIED POLY(3-CAPROLACTONE) NANOPARTICLES 

Poly(3-caprolactone) (PCL) is another biodegradable polymer that has been used to encapsulate 

hydrophobic drugs. Using this polymer, nanoparticles were prepared by solvent displacement in an 

acetone-water system in the presence of Pluronic w . The solvent displacement technique used for 

the preparation of PCL nanoparticles facilitates instant adsorption of PPO–PEO groups when the 

organic solution of the polymer is introduced into aqueous solution containing the stabilizer. 51 In 

addition, it also favors the encapsulation of hydrophobic drugs such as tamoxifen that could be 

dissolved along with the polymer in the organic phase, resulting in a high entrapment efficiency of 

greater than 90% at loading levels of 20% of the weight of the drug. The intracellular uptake of 

these nanoparticles in MCF-7 estrogen receptor-positive breast cancer cells and MDA-MB231 

human breast adenocarcinoma cells was monitored at different time points using tritiated [ 3 H]tamoxifen. 

The results showed that the cell uptake followed saturable kinetics with most of the 

nanoparticles being internalized within the first 30 min of incubation. 52 

The in vivo disposition of these PEO-modified PCL nanoparticles was completed in mice 

bearing MDA-MB231 xenograft breast cancer tumors as it is a well-characterized and simpler 

model compared to MCF-7 that requires estrogen priming for growth. This study was used to 

compare the biodistribution profiles of the nanoparticles modified by the Pluronic w F-68 containing 

30 residues of propylene oxide (PO) and 76 residues of ethylene oxide (EO) with those of Pluronic w 

F-108 that has 56 PO residues and 122 EO residues. The nanoparticulate formulations encapsulated 

with radiolabled tamoxifen were used. As expected, upon intravenous administration, the Pluronic w - 

modified nanoparticles had higher tumor concentrations in comparision to the tamoxifen solution 

and drug-loaded, unmodified nanoparticles. At early time points (one hour), the nanoparticles 

modified with Pluronic w F-108 had greater concentration in the tumor with no significant difference 

in the concentration of the Pluronic w -modified formulations at six hours post-injection. A similar 

trend has been observed in plasma concentration-time profiles with PEO–PCL formulations circulating 

for a longer time than the controls. 40 

In a parallel study, the PEO-modified PCL nanoparticles were radiolabeled by a similar 

procedure specific to PEO–PBAE nanoparticles. The nanoparticles, encapsulated with [ 3 H] 

tritium-labeled paclitaxel, were used to understand the change in concentration and localization 

of the drug in ovarian tumors (SKOV3). From the biodistribution studies, it was shown that the 

modification of PCL nanoparticles with PEO had extended the mean residence time to up to 25 h. 

Hydrophobic drugs such as paclitaxel were found to have high plasma concentrations as a result of 

their protein binding capacity; however, they were cleared from the blood within 24 h. The 

circulation time of such drugs has been enhanced by encapsulating them in PEO–PCL nanoparticles 

that, in turn, has resulted in higher concentrations of the drug in the tumors. The PEO–PCL 

nanoparticles have resulted in an 8.7-fold increase in drug concentration at five-hour time points 

when compared to the solution form of the drug. 40 

In another study, PEO-modified PCL nanoparticles were used for combination therapy of 

ceramide with paclitaxel in order to overcome MDR, particularly in cases of breast and ovarian 



cancers. Several MDR specimens of cancer were found to exhibit elevated levels of the enzyme 

glucosylceramide synthease (GCS) (also called ceramide glucosyl transferase) that is responsible 

for the inactivation of ceramide, a messenger in apoptotic signaling to its non-functional moiety 

glucosylceramide. 53–55 These findings suggest the importance of the role of ceramide in the 

mediation of a cytotoxic response and its function as an apoptotic messenger in the signaling 

pathway. 

To potentially overcome MDR in ovarian cancer cell lines, C6-ceramide has been encapsulated 

along with paclitaxel into PEO-modified PCL nanoparticles. Upon treatment of the cells with 

paclitaxel, the MDR cell line SKOV3/TR exhibited 65.65G2.16% viability at 1 mM dose; the 

sensitive cell line SKOV3 showed 16.37G0.41% viability at 100 nM dose. Co-treatment of 

these cells along with 20 mM C6-ceramide in addition to paclitaxel (1 mM in the case of a resistant 

cell line and 100 nM in a sensitive cell line) resulted in a cell viability of 2.69G0.51% with the 

resistant cells and 7.38G1.25% with the sensitive cell lines, indicating a significant increase in cell 

death when compared to the paclitaxel treatment alone. Furthermore, the co-encapsulation of these 

drugs within PEO–PCL nanoparticles resulted in enhanced cell kill compared to the drugs alone. 

A 10 nM dose of paclitaxel, delivered in combination with ceramide in PEO–PCL nanoparticles, 

resulted in 63.98G4.9% viability, and the free drugs in solution at these doses did not provoke any 

cell kill in the resistant cell line. The use of these drug-loaded nanoparticles resulted in a 100-fold 

increase in chemosensitivity of the MDR cells. These results demonstrate the clinical use of 

PEO–PCL nanoparticles in overcoming MDR by combination therapy. 

13.7 ILLUSTRATIVE EXAMPLES OF POLYMERIC NANOPARTICLES 

FOR GENE DELIVERY 

13.7.1 POLYETHYLENE GLYCOL-MODIFIED GELATIN NANOPARTICLES 

To develop safe and effective systemically administered non-viral gene therapy vectors for solid 

tumors, PEGylated gelatin was synthesized and fabricated into nanoparticles. The PEG-modified 

gelatin was synthesized by reacting Type-B gelatin with PEG-epoxide that was further confirmed 

by electron spectroscopy for chemical analysis (ESCA), the results of which had shown an increase 

in the presence of an ether carbon (–C–O–) peak in the PEGylated gelatin nanoparticles. The 

experiments performed using these polymeric nanoparticulate carriers proved their ability to encapsulate 

DNA, improved transfection efficiency in vitro, and enhanced the biodistribution profile 

when the carriers were injected into mice bearing tumors. 56–58 

The in vitro cytotoxicity assays indicate that both gelatin and PEGylated gelatin are non-toxic 

to the cells. The cell uptake and trafficking studies of the control (gelatin) and PEGylated gelatin 

nanoparticles encapsulated with electron-dense gold nanoparticles confirmed that the particles were 

internalized by an endocytic pathway and remained stable during the vesicular transport. Further, 

the transfection studies carried out using the DNA (pEGFP-N1) encapsulated nanoparticles were 

internalized in NIH-3T3 fibroblast cells within the first six hours of incubation. Green fluorescent 

protein expression was observed after 12 h of nanoparticle incubation and remained stable for up to 

96 hours. Flow cytometry results showed that the DNA transfection efficiency of PEGylated gelatin 

nanoparticles was better than that of gelatin. 57 

In order to determine the biodistribution profile, nanoparticles were radiolabeled with iodine- 

125 [ 125 I] and intravenously injected into mice bearing Lewis lung carcinoma (LLC) tumors. From 

the radioactivity of the plasma and other organs collected, it was evident that the majority of the 

PEGylated gelatin nanoparticles remained in the blood pool or were taken up by the tumor mass and 

liver. A two-fold increase in concentration of the PEGylated nanoparticles was observed in plasma 

even at three-hour time points, and about 4–5% of the recovered dose of PEGylated gelatin 

nanoparticles was found to be present in the tumor mass for up to 12 h. Upon non-compartmental 

pharmacokinetic analysis of the plasma and tumor concentration-time profiles, it was observed 


238 

that PEGylated gelatin nanoparticles had greater mean residence time in plasma and a more than 

six-fold increase in half-life in the tumor. The results of this study showed the stealth nature of 

the PEGylated gelatin nanoparticles in avoiding uptake by RES, allowing the nanoparticles to 

circulate longer in plasma. 59 

TheinvivotransfectionefficiencyofplasmidDNA-encapsulated(pCMV-b encoding 

b-galactosidase) PEGylated gelatin nanoparticles was evaluated by injecting these nanoparticles 

intravenously into LLC tumor-bearing female C57BL/6J mice. Following systemic administration, 

the animals were sacrificed, and the transgene expression in different organs was quantitatively 

determined by using o-nitrophenyl-b-d-galactopyranosidase (ONPG), a clear substrate of 

b-galactosidase that is converted to a yellow-colored product with an absorbance maximum at 

420 nm, and qualitatively by X-gal w tissue staining. When compared to the gelatin nanoparticles, 

the PEGylated gelatin nanoparticles were found to efficiently transfect the tumor cells with the 

b-galactosidase expression increasing at time points up to 96 h post-transfection with the absorbance 

value at 420 nm increasing from 0.6 to 0.85 starting from 12 h post-transfection. The results 

of these studies clearly indicate that the long circulating, biocompatible, and biodegradable nanoparticles 

of PEGylated gelatin would be ideal for gene delivery applications to solid tumors. 60 

13.7.2 POLYETHYLENE GLYCOL-MODIFIED THIOLATED-GELATIN NANOPARTICLES 

Another study synthesized thiolated gelatin and formulated it into nanoparticles that could rapidly 

release their payload in the presence of a high concentration of glutathione. The thiolated gelatins 

with different degrees of thiolation were characterized by Ellman’s assay to determine degree of 

thiolation and by in vitro cytotoxicity assay. Upon carrying out the qualitative and quantitative 

transfection studies, the nanoparticles of thiolated gelatin, made with 20 mg of aminothiolane per 

gram of gelatin (SHGel-20) and encapsulated with plasmid DNA (pEGFP-N1), were found to have 

higher transfection efficiency in mouse fibroblast (NIH-3T3) cells in vitro (Figure 13.3). 61 

These nanoparticles, prepared by the desolvation method, were post-PEGylated with polyethylene 

glycol-succinimidyl glutarate (PEG-SG) of molecular weight 2,000 Da. The radiolabeled 

( 111 Indium) long-circulating (PEG-modified) thiolated-gelatin nanoparticles were injected intravenously 

into mice bearing human breast cancer (MDA-MB435) xenografts. The biodistribution and 

pharmacokinetics of these particles were compared to those of nanoparticles prepared from thiolated-gelatin, 

gelatin, and PEG-modified gelatin. The unmodified gelatin and thiolated gelatin 

nanoparticles were shown to have a plasma half-life of approximately three hours. Upon modification 

with PEG, the nanoparticles were found to have prolonged circulation times extending up to 

10.7 h in the case of PEG-gelatin and 15.3 h with PEG-thiolated-gelatin nanoparticles. The 

successful surface hydrophilization of the gelatin and thiolated gelatin nanoparticles resulted in 

enhanced circulation times with almost 7% of the dose remaining in plasma at 12 h. The results also 

showed that the thiolated-gelatin nanoparticles (both non-modified and PEG-modified) had greater 

accumulations in the tumor than the nanoparticles prepared from non-thiolated gelatin. These 

thiolated-gelatin nanoparticles that are an improvement over gelatin nanoparticles can be made 

long-circulating by surface modification with PEG, and they can be used for rapid delivery of the 

payload the tumor tissues. The mechanisms of long-circulation through steric hindrance when in 

circulation and preferential accumulation in the tumor mass or inside the cells as a result of the cleavage 

of the disulfide bonds (arising from thiolation) in the highly reductive tumor-microenvironment 

can be exploited for the delivery of drugs and genes in clinical applications. 40,61 

13.8 CONCLUDING REMARKS 


There are multiple factors affecting the delivery of drugs and genes to tumors. Factors such as blood 

flow, angiogenesis, microvessel density, interstitial pressure, macrophage activity, extracellular and 

intracellular components, and, most importantly, the physicochemical properties of the drug carrier 



play an important role in the transport of drugs and macromolecules to tumors. This lab is particularly 

interested in enhancing the transport of drugs and genes to tumors by altering the 

physicochemical properties of the polymeric carrier used to encapsulate the drugs and genes. As 

shown in the illustrative examples, the biodistribution properties of the polymeric carriers could be 

modified using hydrophilic polymers such as PEG and PEO. This therapeutic strategy could be used 

to alter the passive/active targeting ability of the drug and gene carriers. However, the delivery of 

these newer agents is still a challenge, highlighting the necessity of additional research in this area. 


These studies have been supported by grants R01-CA095522 and R01-CA119617 from the 

National Cancer Institute, National Institutes of Health. Our research work on poly(beta-amino 

ester) nanoparticles is done in collaboration with Professor Robert Langer’s lab at MIT 

(Cambridge, MA). We deeply appreciate the on-going, highly productive collaborations with 

his group and especially the assistance of Dr. Steven Little, Dr. Daniel Anderson, and 

Mr. David Nguyen. 

REFERENCES 

(a) 

FIGURE 13.3 Transfection images of mouse fibroblast (NIH3T3) cells treated with thiolated-gelatin nanoparticles 

encapsulated with plasmid DNA (pEGFP-N1). The images show fluorescence of the expressed green 

fluorescent protein at 6 h (a) and 96 h post-transfection (b). 

(b) 

1. Boehm, T., Folkman, J., Browder, T., and O’Reilly, M. S., Antiangiogenic therapy of experimental 

cancer does not induce acquired drug resistance, Nature, 390(6658), 404–407, 1997. 

2. Kerbel, R. S., A cancer therapy resistant to resistance, Nature, 390(6658), 335–336, 1997. 

3. Kirsch, M., Strasser, J., Allende, R., Bello, L., Zhang, J., and Black, P. M., Angiostatin suppresses 

malignant glioma growth in vivo, Cancer Research, 58(20), 4654–4659, 1998. 

4. Ganjavi, H., Gee, M., Narendran, A., Freedman, M. H., and Malkin, D., Adenovirus-mediated p53 

gene therapy in pediatric soft-tissue sarcoma cell lines: Sensitization to cisplatin and doxorubicin, 

Cancer Gene Therapy, 12(4), 397–406, 2005. 

5. Liu, Y., Huang, H., Saxena, A., and Xiang, J., Intratumoral coinjection of two adenoviral vectors 

expressing functional interleukin-18 and inducible protein-10, respectively, synergizes to facilitate 

regression of established tumors, Cancer Gene Therapy, 9(6), 533–542, 2002. 

6. Van Meir, E. G., Polverini, P. J., Chazin, V. R., Su Huang, H. J., de Tribolet, N., and Cavenee, W. K., 

Release of an inhibitor of angiogenesis upon induction of wild type p53 expression in glioblastoma 

cells, Nature Genetics, 8(2), 171–176, 1994. 

7. Schiffelers, R. M., Ansari, A., Xu, J., Zhou, Q., Tang, Q., Storm, G., Molema, G., Lu, P. Y., Scaria, P. 

V., Woodle, M. C. Cancer siRNA therapy by tumor selective delivery with ligand-targeted sterically 

stabilized nanoparticle Nucleic Acids Research, 32(19), e149, 2004 (http://nar.oxfordjournal.org/cgi/ 

content/abstract/32/19/e149). 


240 


8. Yuan, F., Salehi, H. A., Boucher, Y., Vasthare, U. S., Tuma, R. F., and Jain, R. K., Vascular 

permeability and microcirculation of gliomas and mammary carcinomas transplanted in rat and 

mouse cranial windows, Cancer Research, 54(17), 4564–4568, 1994. 

9. Yuan, F., Leunig, M., Huang, S. K., Berk, D. A., Papahadjopoulos, D., and Jain, R. K., Microvascular 

permeability and interstitial penetration of sterically stabilized (stealth) liposomes in a human tumor 

xenograft, Cancer Research, 54(13), 3352–3356, 1994. 

10. Jain, R. K., Physiological barriers to delivery of monoclonal antibodies and other macromolecules in 

tumors, Cancer Research, 50(Suppl. 3), 814s–819s, 1990. 

11. Jang, S. H., Wientjes, M. G., Lu, D., and Au, J. L., Drug delivery and transport to solid tumors, 

Pharmaceutical Research, 20(9), 1337–1350, 2003. 

12. Maeda, H., SMANCS and polymer-conjugated macromolecular drugs: Advantages in cancer 

chemotherapy, Advanced Drug Delivery Reviews, 46(1–3), 169–185, 2001. 


Critical Reviews in Therapeutic Drug Carrier Systems, 6(3), 193–210, 1989. 


chemotherapy: Mechanism of tumoritropic accumulation of proteins and the anti-tumor agent 

smancs, Cancer Research, 46(12 Pt 1), 6387–6392, 1986. 

15. Balkwill, F. and Mantovani, A., Inflammation and cancer: Back to Virchow?, Lancet, 357(9255), 

539–545, 2001. 

16. Coussens, L. M. and Werb, Z., Inflammation and cancer, Nature, 420(6917), 860–867, 2002. 

17. Lewis, C. and Murdoch, C., Macrophage responses to hypoxia: Implications for tumor progression and 

anti-cancer therapies, American Journal of Pathology, 167(3), 627–635, 2005. 

18. Mantovani, A., Sozzani, S., Locati, M., Allavena, P., and Sica, A., Macrophage polarization: Tumorassociated 

macrophages as a paradigm for polarized M2 mononuclear phagocytes, Trends in 

Immunology, 23(11), 549–555, 2002. 

19. Luo, D., Haverstick, K., Belcheva, N., Han, E., and Saltzman, M., Poly(ethylene glycol)-conjugated 

PAMAM dendrimer for biocompatible, high-efficiency DNA delivery, Macromolecules, 35, 

3456–3462, 2002. 

20. Kircheis, R., Blessing, T., Brunner, S., Wightman, L., and Wagner, E., Tumor targeting with surfaceshielded 

ligand—polycation DNA complexes, Journal of Controlled Release, 72(1–3), 165–170, 

2001. 

21. Kim, S. H., Jeong, J. H., Cho, K. C., Kim, S. W., and Park, T. G., Target-specific gene silencing by 

siRNA plasmid DNA complexed with folate-modified poly(ethylenimine), Journal of Controlled 

Release, 104, 223–232, 2005. 

22. Suh, W., Han, S. O., Yu, L., and Kim, S. W., An angiogenic, endothelial-cell-targeted polymeric gene 

carrier, Molecular Therapy, 6(5), 664–672, 2002. 

23. Kursa, M., Walker, G. F., Roessler, V., Ogris, M., Roedl, W., Kircheis, R., and Wagner, E., Novel 

shielded transferrin-polyethylene glycol-polyethylenimine/DNA complexes for systemic tumortargeted 

gene transfer, Bioconjugate Chemistry, 14(1), 222–231, 2003. 

24. Lee, H., Kim, T. H., and Park, T. G., A receptor-mediated gene delivery system using streptavidin and 

biotin-derivatized, pegylated epidermal growth factor, Journal of Controlled Release, 83(1), 109–119, 

2002. 

25. Merdan, T., Callahan, J., Petersen, H., Kunath, K., Bakowsky, U., Kopeckova, P., Kissel, T., and 

Kopecek, J., Pegylated polyethylenimine-Fab’ antibody fragment conjugates for targeted gene 

delivery to human ovarian carcinoma cells, Bioconjugate Chemistry, 14(5), 989–996, 2003. 

26. Natarajan, A., Xiong, C. Y., Albrecht, H., DeNardo, G. L., and DeNardo, S. J., Characterization of 

site-specific ScFv PEGylation for tumor-targeting pharmaceuticals, Bioconjugate Chemistry, 16(1), 

113–121, 2005. 


Advanced Drug Delivery Reviews, 56(11), 1649–1659, 2004. 

28. Ehrlich, P., Studies in Immunity, Wiley, New York, 1906. 

29. Ringsdorf, H., Structure and properties of pharmacologically active polymers, Journal of Polymer 

Science Polymer Symposium, 51, 135–153, 1975. 



30. Stastny, M., Strohalm, J., Plocova, D., Ulbrich, K., and Rihova, B., A possibility to overcome 

p-glycoprotein (PGP)-mediated multidrug resistance by antibody-targeted drugs conjugated to 

N-(2-hydroxypropyl)methacrylamide (HPMA) copolymer carrier, European Journal of Cancer, 35, 

459–466, 1999. 

31. El-Aneed, A., An overview of current delivery systems in cancer gene therapy, Journal of Controlled 

Release, 94(1), 1–14, 2004. 

32. Fenske, D. B., MacLachlan, I., and Cullis, P. R., Long-circulating vectors for the systemic delivery of 

genes, Current Opinion in Molecular Therapeutics, 3(2), 153–158, 2001. 

33. Han, S., Mahato, R. I., Sung, Y. K., and Kim, S. W., Development of biomaterials for gene therapy, 

Moleular Therapeutics, 2(4), 302–317, 2000. 

34. McCormick, F., Cancer gene therapy: Fringe or cutting edge?, Nature Reviews Cancer, 1(2), 

130–141, 2001. 

35. Lehrman, S., Virus treatment questioned after gene therapy death, Nature, 401(6753), 517–518, 1999. 

36. Marshall, E., Gene therapy. Second child in French trial is found to have leukemia, Science, 

299(5605), 320, 2003. 

37. Thomas, C. E., Ehrhardt, A., and Kay, M. A., Progress and problems with the use of viral vectors for 

gene therapy, Nature Reviews Genetics, 4(5), 346–358, 2003. 

38. Kommareddy, S., Tiwari, S. B., and Amiji, M. M., Long-circulating polymeric nanovectors for 

tumor-selective gene delivery, Technology in Cancer Research and Treatment, 4(6), 615–626, 2005. 

39. Torchilin, V. P., Drug targeting, European Journal of Pharmaceutial Sciences, 11(Suppl. 2), 

S81–S91, 2000. 

40. Kommareddy, S., Shenoy, D. B., and Amiji, M., Nanoparticulate carriers of gelatin and gelatin 

derivatives, In Biological and Pharmaceutical Nanomaterials, Kumar, C., Ed., Wiley-VCH, Weinheim, 

Germany, pp. 330–347, 2005. 

41. Torchilin, V. P., Polymer-coated long-circulating microparticulate pharmaceuticals, Journal of 

Microencapsulation, 15(1), 1–19, 1998. 

42. Torchilin, V. P. and Papisov, M. I., Why do polyethylene glycol-coated liposomes circulate so long?, 

Journal of Liposome Research, 4(1), 725–739, 1994. 

43. Zalipsky, S., Functionalized Poly(ethylene glycol) for preparation of biologically revalant conjugates, 


44. Gref, R., Domb, A. J., Quellec, P., Blunk, T., Müller, R. H., Verbavatz, J. M., and Langer, R., The 

controlled intravenous delivery of drugs using PEG-coated sterically stabilized nanospheres, 

Advanced Drug Delivery Reviews, 16(2–3), 215–233, 1995. 

45. Delgado, C., Francis, G. E., and Fisher, D., The uses and properties of PEG-linked proteins, Critical 

Reviews in Therapeutic Drug Carrier Systems, 9(3–4), 249–304, 1992. 

46. Mehvar, R., Dextrans for targeted and sustained delivery of therapeutic and imaging agents, Journal 

of Controlled Release, 69(1), 1–25, 2000. 

47. Torchilin, V. P., How do polymers prolong circulation time of liposomes?, Journal of Liposome 

Research, 6(1), 99–116, 1996. 

48. Oupicky, D., Howard, K. A., Konak, C., Dash, P. R., Ulbrich, K., and Seymour, L. W., Steric 

stabilization of poly-L-Lysine/DNA complexes by the covalent attachment of semitelechelic 

poly[N-(2-hydroxypropyl)methacrylamide], Bioconjugate Chemistry, 11(4), 492–501, 2000. 

49. Lynn, D. M., Amiji, M. M., and Langer, R., pH-responsive polymer microspheres: Rapid release of 

encapsulated material within the range of intracellular pH, Angewandte Chemie (International ed. in 

English), 40(9), 1707–1710, 2001. 

50. Potineni, A., Lynn, D. M., Langer, R., and Amiji, M. M., Poly(ethylene oxide)-modified 

poly(beta-amino ester) nanoparticles as a pH-sensitive biodegradable system for paclitaxel delivery, 

Journal of Controlled Release, 86(2–3), 223–234, 2003. 

51. Scholes, P. D., Coombes, A. G., Illum, L., Davis, S. S., Watts, J. F., Ustariz, C., Vert, M., and Davies, 

M. C., Detection and determination of surface levels of poloxamer and PVA surfactant on biodegradable 

nanospheres using SSIMS and XPS, Journal of Controlled Release, 59(3), 261–278, 1999. 

52. Chawla, J. S. and Amiji, M. M., Cellular uptake and concentrations of tamoxifen upon administration 

in poly(epsilon-caprolactone) nanoparticles, AAPS PharmSci, 5(1), E3, 2003 (http://www.aapsj.org/ 

view.asp?art=ps050103). 


242 


53. Lavie, Y., Cao, H., Bursten, S. L., Giuliano, A. E., and Cabot, M. C., Accumulation of glucosylceramides 

in multidrug-resistant cancer cells, Journal of Biological Chemistry, 271, 19530–19536, 1996. 

54. Liu, Y. Y., Han, T. Y., Giuliano, A. E., and Cabot, M. C., Ceramide glycosylation potentiates cellular 

multidrug resistance, FASEB Journal, 15(3), 719–730, 2001. 

55. Morjani, H., Aouali, N., Belhoussine, R., Veldman, R. J., Levade, T., and Manfait, M., Elevation of 

glucosylceramide in multidrug-resistant cancer cells and accumulation in cytoplasmic droplets, 

International Journal of Cancer, 94, 157–165, 2001. 

56. Kaul, G. and Amiji, M., Long-circulating poly(ethylene glycol)-modified gelatin nanoparticles for 

intracellular delivery, Pharmaceutical Research, 19(7), 1062–1068, 2002. 

57. Kaul, G. and Amiji, M., Cellular interactions and in vitro DNA transfection studies with poly(ethylene 

glycol)-modified gelatin nanoparticles, Journal of Pharmaceutical Sciences, 94(1), 184–198, 2004. 

58. Kaul, G., Lee-Parsons, C., and Amiji, M., Poly(ethylene glycol)-modified gelatin nanoparticles for 

intracellular delivery, Pharmaceutical Engineers, 23(5), 1–5, 2003. 

59. Kaul, G. and Amiji, M., Biodistribution and targeting potential of poly(ethylene glycol)-modified 

gelatin nanoparticles in subcutaneous murine tumor model, Journal of Drug Targeting, 12(9–10), 

585–591, 2004. 

60. Kaul, G. and Amiji, M., Tumor-targeted gene delivery using poly(ethylene glycol)-modified gelatin 

nanoparticles: In vitro and in vivo studies, Pharmaceutical Research, 22(6), 951–961, 2005. 

61. Kommareddy, S. and Amiji, M., Preparation and evaluation of thiol-modified gelatin nanoparticles for 

intracellular DNA delivery in response to glutathione, Bioconjugate Chemistry, 16(6), 1423–1432, 

2005. 


14 

CONTENTS 

Biodegradable PLGA/PLA 

Nanoparticles for Anti-Cancer 

Therapy 

Sanjeeb K. Sahoo and Vinod Labhasetwar 

14.1 Introduction ....................................................................................................................... 243 

14.2 General Principles of Drug Targeting to Cancer.............................................................. 244 

14.2.1 Passive Targeting ................................................................................................ 244 

14.2.2 Active Targeting.................................................................................................. 244 

14.3 PLGA as a Polymer for Nanoparticles ............................................................................. 244 

14.4 Application of PLGA/PLA Nanoparticles as Drug Delivery Vehicles 

to Cancer Tissues .............................................................................................................. 245 

14.5 PLGA Nanoparticles for Gene Delivery to Cancer.......................................................... 247 

14.6 PLGA Nanoparticles for Photodynamic Therapy............................................................. 248 



References..................................................................................................................................... 249 


Despite significant efforts in the field of oncology, cancer remains one of the leading causes of death 

in industrialized countries. 1 Of the various options available, surgery and radiation therapy are 

commonly used to treat localized tumors whereas chemotherapy is primarily used in the management 

and elimination of hematological malignancies and metastasized tumors. 2 The major 

limitation of the current cancer chemotherapeutics is their high pharmacokinetic volume of distribution 

and rapid elimination rates, requiring more frequent and high dose administration of these 

agents. 3 Therefore, cancer chemotherapeutics cause unacceptable damage to normal tissue when 

used at doses required to eradicate cancer cells. Furthermore, multidrug resistance in tumor cells 

exacerbates the problem of cancer chemotherapy. 4–7 Selective delivery of anti-cancer drugs to the 

tumor tissue offers a formidable solution to the problem. This could significantly enhance the 

therapeutic efficacy of these drugs and diminish their toxicity. 8,9 The overall goal of drug delivery 

strategies is to selectively attack cancer cells while saving normal tissue from drug toxicity. 10,11 

The other issue with cancer chemotherapeutics is their systemic delivery, as most of these agents 

are poorly water soluble. Hence they require adjuvant or excipients to dissolve, which are often 

associated with serious side effects. 12 Therefore, drug carrier systems which would address the 


243

244 

above problem of delivery as well as target drugs selectively to the tumor would be a major 

advancement in cancer chemotherapy. Many macromolecular carriers, including soluble synthetic 

and natural polymers, implants, liposomes, microspheres, dendrimers, nanoparticles, etc., have 

been tested for selective delivery of drugs to the tumor tissue. 13 

14.2 GENERAL PRINCIPLES OF DRUG TARGETING TO CANCER 

14.2.1 PASSIVE TARGETING 

Passive targeting refers to the accumulation of drug or drug-carrier system at a particular site due to 

physicochemical or pharmacological factors. Permeability of the tumor vasculature increases to the 

point where particulate carriers such as nanoparticles can extravasate from blood circulation and 

localize in the tumor tissue. 14,15 This occurs because as tumors grow and begin to outstrip the 

available supply of oxygen and nutrients, they release cytokines and other signaling molecules that 

recruit new blood vessels to the tumor, a process known as angiogenesis. 16 Angiogenic blood 

vessels, unlike the tight blood vessels in most normal tissues, have gaps as large as 600–800 nm 

between adjacent endothelial cells. Drug carriers in the nanometer size range can extravasate 

through these gaps into the tumor interstitial space. 17,18 Because tumors have impaired lymphatic 

drainage, the carriers concentrate in the tumor, resulting in higher drug concentration in the tumor 

tissue (10-fold or higher) than that can be achieved with the same dose of free drug. This is 

commonly referred to as enhanced permeability and retention, or the EPR effect. 

14.2.2 ACTIVE TARGETING 

Active targeting to the tumor can be achieved by molecular recognition of cancer cells either via 

ligand–receptor or antibody–antigen interactions. Active targeting may also lead to receptormediated 

cell internalization of drug carrier system. Nanoparticles and other polymer drugconjugates 

offer numerous opportunities for targeting tumors through surface modifications 

which allow specific biochemical interactions with the proteins/receptorsexpressedontarget 

cells. 19,20 For active and passive targeting of drug carrier systems, it is essential to avoid their 

uptake by the reticuloendothelial system (RES) so that they remain in the blood circulation and 

extravasate in the tumor vasculature. Particles with more hydrophobic surfaces are preferentially 

taken up by the liver, followed by the spleen and lungs. 21,22 Size of nanoparticles as well as their 

surface characteristics are the key parameters that can alter the biodistribution of nanoparticles. 

Particles smaller than 100 nm and coated with hydrophilic polymers such as amphiphilic polymeric 

compounds which are made of polyethylene oxide such as poloxamers, poloxamines, or polyethylene 

glycol (PEG) are being investigated to avoid their uptake by the RES. To improve the efficacy 

of targeting cancer chemotherapeutics to the tumor, a combination of passive and active targeting 

strategy is being investigated where long-circulating drug carriers are conjugated to tumor cellspecific 

antibody or peptides. 23 In addition to the above approach, direct intratumoral injection of 

the carrier system is feasible if the tumor is localized and can be accessed for administration of a 

carrier system. 24 

14.3 PLGA AS A POLYMER FOR NANOPARTICLES 


A number of different polymers, both synthetic and natural, have been utilized in formulating 

biodegradable nanoparticles. Synthetic polymers have the advantage of sustaining the release of 

the encapsulated therapeutic agent over a period of days to several weeks as compared to natural 

polymers which, in general, have a relatively short duration of drug release. The polymers used for 

the formulation of nanoparticles include synthetic polymers such as polylactide–polyglycolide 

copolymers, polyacrylates, and polycaprolactones, or natural polymers such as albumin, gelatin, 


Biodegradable PLGA/PLA Nanoparticles for Anti-Cancer Therapy 245 

alginate, collagen, and chitosan. Of these polymers, polylactides (PLA) and poly (D,L-lactide-coglycolide) 

(PLGA) have been most extensively investigated for drug delivery applications. 19 

PLGA/PLA-based polymers have a number of advantages over other polymers used in drug and 

gene delivery, such as their biodegradability, biocompatibility, and approval by the FDA for human 

use. 19,25 PLGA/PLA polymers degrade in the body through hydrolytic cleavage of the ester linkage 

to lactic and glycolic acid, although there are reports of involvement of enzymes in their biodegradation. 

19,25 These monomers are easily metabolized in the body via Krebs’ cycle and eliminated 

as carbon dioxide and water. Biodegradation products of PLGA and PLA polymers are formed at a 

very slow rate, and they therefore do not affect normal cell function. 26 Furthermore, these polymers 

have been tested for toxicity and safety in extensive animal studies and are currently used in 

humans for resorbable sutures, bone implants and screws, contraceptive implants, 27,28 and also 

as graft materials for artificial organs and supporting scaffolds in tissue engineering research. 28–30 

Long-term biocompatibility of these polymers was demonstrated by the absence of any untoward 

effects on intravascular administration of nanoparticles formulated using these polymers to the 

arterial tissue in pig and rat models of restenosis. 31 

14.4 APPLICATION OF PLGA/PLA NANOPARTICLES AS DRUG DELIVERY 

VEHICLES TO CANCER TISSUES 

There are several studies regarding PLGA/PLA nanoparticles or some modification of these polymers 

for delivery of anti-cancer agents and other therapeutic agents. 19 We have recently 

demonstrated increased efficacy of transferrin conjugated paclitaxel-loaded PLGA nanoparticles 

(Figure 14.1a and b) both in vitro and in an animal model of prostate carcinoma. 24 Transferrin 

receptors are over-expressed in most cancer cells by two to tenfold more than in normal cells. We 

have demonstrated that transferring-conjugated nanoparticles have enhanced cellular uptake 

(Figure 14.1c) and retention than unconjugated nanoparticles. 4 A single-dose intratumoral injection 

of transferrin conjugated nanoparticles in animal models of prostate carcinoma demonstrated 

complete tumor regression and higher survival rate than animals that received either drug in 

solution or unconjugated nanoparticles. 24 The IC50 for paclitaxel with transferrin conjugated nanoparticles 

was fivefold lower than that with unconjugated nanoparticles or with drug in solution in 

PC3 24 (Figure 14.1d) and in MCF-7 cells. 4 Kim et al. have demonstrated enhanced intracellular 

delivery of PLGA nanoparticles, which were surface-coated with cationic di-block copolymer, 

poly(L-lysine)–poly(ethylene glycol)–folate (PLL–PEG–FOL), in KB cells that overexpress 

folate receptors. 32 In another study, paclitaxel-loaded PLGA nanoparticles, which were conjugated 

to wheat germ agglutinin (WGA), demonstrated greater anti-proliferation activity in A549 and 

H1299 cells as compared to the conventional paclitaxel formulations. This enhanced activity of 

WGA-conjugated nanoparticles was attributed to greater intracellular accumulation of drug via 

WGA-receptor-mediated endocytosis of conjugated nanoparticles. 33 Cegnar et al. have developed 

cystatin-loaded PLGA nanoparticles with the strategy of inhibiting the tumor-associated activity of 

intracellular cysteine proteases cathepsins B and L. In an in vitro study, cystatin-loaded PLGA 

nanoparticles demonstrated 160-fold greater cytotoxic effect in MCF-10A neoT cells than free 

cystatin. 34 Similarly, interferon-alpha (IFN-alpha) loaded PLGA nanoparticles are being developed 

to improve the therapeutic efficacy of IFN-alpha while reducing its dose-related side effects. 35 In 

studies by Yoo et al., doxorubicin was chemically conjugated to a terminal end group of PLGA by 

an ester linkage and the doxorubicin–PLGA conjugate was formulated into nanoparticles. Nanoparticles 

containing the conjugate exhibited sustained doxorubicin release profiles over a onemonth 

period, whereas those containing unconjugated doxorubicin showed a rapid drug release 

within five days. The conjugated doxorubicin nanoparticles demonstrated increased drug uptake in 

HepG2 cell line and exhibited slightly lower IC 50 value compared to that of free doxorubicin. 


246 

Uptake (μg/mg cell protein) 

(a) 

(c) 

50 

40 

30 

20 

10 

0 

NP NP-Tf NP-Tf+freeTf 

Treatments 

Cumulative % release 

% Growth 

0 

0 10 20 30 40 50 60 70 

The in vivo anti-tumor activity assay showed that a single injection of these nanoparticles had 

comparable activity to that of free doxorubicin when administered by daily injection. 36 Thus 

different strategies are being investigated using biodegradable PLGA-based nanoparticles to 

deliver anti-cancer drugs more effectively, both at the cellular level and also to the tumor tissue. 

60 

50 

40 

30 

20 

10 

(b) 

90 

80 

70 

60 

50 

40 

30 

20 

10 

0 

(d) 


Time (Day) 

1 10 100 1000 

Paclitaxel concentration (ng/ml) 

FIGURE 14.1 (a) Transmission electron micrographs of paclitaxel-loaded PLA nanoparticles (NPs) (barZ 

100 nm). (b) Cumulative release of paclitaxel (Tx) from NPs in vitro. Tx loading in NPs is 5.4% (w/w). Data as 

mean (Gs.e.m. (nZ3). (c) Uptake of unconjugated (NPs) and conjugated (NPs–Tf) in PC3 cells. A suspension 

of NPs (100 mg/mL) was incubated with PC3 cells (5!10 4 cells) for 1 h, cells were washed, and the NP levels 

in cells were determined by HPLC. To determine the competitive inhibition of uptake of Tf-conjugated NPs, 

excess dose of free Tf (50 mg) was added in the medium prior to incubation with NPs–Tf. Data as mean G 

s.e.m. (nZ6).* p!.05 NPs–TfC free Tf vs. NPs, ** p!.005 NPs–Tf vs. NPs. NPs contain a fluorescent 

marker, 6-coumarin, to quantify their uptake. (d) Dose-dependent cytotoxicity of paclitaxel (Tx) in PC3 cells. 

Different concentration of Tx either as solution (%) or encapsulated in NPs (Tx–NPs (&) or Tx–NPs–Tf, (:) 

was added to wells with NPs (without drug) or medium acting as respective controls. The medium was changed 

at two days and then on every alternate day with no further dose of the drug added. The extent of growth 

inhibition was measured at five days. The percentage of survival was determined by standardizing the absorbance 

of controls to 100% (nZ6). Data as mean Gs.e.m. * p!.005 Tx–NPs–Tf vs. Tx–sol and Tx–NPs. (Data 

reproduced from Sahoo, S. K. and Labhasetwar, V., Mol. Pharm., 2, 373, 2005; Sahoo, S. K., Ma, W., and 

Labhasetwar, V., Int. J. Cancer, 112, 335, 2004. With permission.) 



14.5 PLGA NANOPARTICLES FOR GENE DELIVERY TO CANCER 

Gene therapy to cancer represents one of the most rapidly developing areas in pre-clinical and 

clinical cancer research. Crucial to the success of any gene therapy strategy is the efficiency with 

which the gene is delivered. 37 This in turn is dependent upon the type of delivery vector used. Many 

vectors have been developed based on either recombinant viruses or non-viral vectors. Research 

utilizing viral vectors has progressed much more rapidly than that with non-viral vectors. This is 

reflected in the fact that approximately 85% of current clinical protocols involve viral vectors. 

However, these vectors are still limited in many ways, particularly in relation to issues of safety, 

immunogenicity, limitations on the size of the gene that can be delivered, specificity, production 

problems, toxicity, cost and others. 38,39 

Although various polymeric systems are under investigation, 40 our efforts are focused on 

investigating biodegradable nanoparticles as a gene delivery system in cancer therapy. 19,41–43 

These nanoparticles are formulated using PLGA and PLA polymers, with plasmid DNA (pDNA) 

entrapped into the nanoparticle matrix. The main advantage of PLGA or PLA-based nanoparticles 

for gene delivery is their non-toxic nature. Furthermore, the slow release of the encapsulated DNA 

from nanoparticles is expected to provide sustained gene delivery (Figure 14.2a). Blends of PLGA 

and polyoxyethylene derivatives such as poloxamer and poloxamine are also being used to 

modulate DNA release from nanoparticles. 44 In our study we demonstrated anti-proliferative 

activity of wild-type (wt) p53 gene-loaded nanoparticles in breast cancer cell line. 45 Cells transfected 

with wt-p53 DNA-loaded nanoparticles demonstrated sustained and significantly greater 

anti-proliferative effect than those treated with naked wt-p53 DNA (Figure 14.2b) or wt-p53 

DNA complexed with a commercially available transfecting agent (Lipofectamine). This effect 

was attributed to sustained nanoparticles-mediated gene expression. This was evident from 

sustained p53 mRNA levels observed in cells transfected with nanoparticles compared with 

levels in cells which were transfected with naked wt-p53 DNA. 

Fluorescence intensity 

1.2 

1 

0.8 

0.6 

0.4 

0.2 

Nanoparticles 

DNA 

% Proliferation 

0 

0.08 1 3 5 7 

(a) Time (day) 

(b) 

120 

100 

80 

60 

40 

20 

0 

p53 DNA 

p53-Nanoparticles 

0 2 4 6 8 

Time (day) 

FIGURE 14.2 (a) Quantitative determination of intracellular DNA levels. Cells transfected with YOYOlabeled 

DNA-loaded nanoparticles demonstrated sustained and increase in intracellular DNA levels as 

opposed to transient DNA levels in the cells transfected with naked DNA. Data represented as mean G 

s.e.m., nZ6, * p!.001. (b) Anti-proliferative activity of wt-p53 DNA-loaded nanoparticles (NPs) and 

naked wt-p53 DNA (DNA) in MDAMB-435S cells. Cells (2500 cells/well) grown in a 96-well plate were 

incubated with DNA-loaded NPs or equivalent amount of naked DNA Cell growth was followed using a 

standard MTS assay. NPs demonstrated increase in anti-proliferative activity with incubation time. Data 

represented as meanGs.e.m., nZ6, * p!.01. (Data reproduced from Prabha, S. and Labhasetwar, V., Mol. 

Pharm., 1, 211, 2004. With permission.) 


248 

Recently, He et al. have formulated thymidine kinase gene (TK gene)-loaded nanoparticles and 

investigated the expression of TK gene in hepatocarcinoma cells. 46 The expression of DNA 

encapsulated in nanoparticles was significantly greater than that with naked DNA. In another 

study, Cohen et al. have demonstrated significantly greater gene transfection (1–2 orders of magnitude) 

with PLGA nanoparticles than that with a liposomal formulation following their 

intramuscular injection in rats. Furthermore, the nanoparticle-mediated gene transfection was 

seen up to 28 days in the above study. Recently Kumar et al. have prepared cationic PLGA 

nanoparticles modified with cationic chitosan and demonstrated gene expression in A549 epithelial 

cells. 47 

14.6 PLGA NANOPARTICLES FOR PHOTODYNAMIC THERAPY 

Photodynamic therapy (PDT) is a promising new treatment modality 48 which involves administration 

of a photosensitizing drug. 49 Upon illumination at a suitable wavelength, it induces a 

photochemical reaction resulting in generation of reactive oxygen species (ROS) that can kill 

tumor cells directly as well as the tumor-associated vasculature, leading to tumor infarction. 

Targeting is essential in PDT as singlet oxygen is extremely reactive. Several strategies such as 

chemical conjugation with various water-soluble polymers and encapsulation into colloidal carriers 

have been considered for the parenteral administration of photosensitizers. These colloidal carriers 

include liposomes, emulsions, polymeric micelles and nanoparticles. PLGA nanoparticles were 

evaluated for PDT using meso-tetra(4-hydroxyphenyl) porphyrin as photosensitizer. 50 Vargas et al. 

have encapsulated porphyrin in PLGA nanoparticles and demonstrated enhanced photodynamic 

activity against mammary tumor cells than free drug. 51 Colloidal carrier associated with photosensitizer 

showed enhanced photodynamic efficiency and selectivity of tumor targeting as 

compared with dye administered in homogenous aqueous solution. 52,53 Recently, Saxena et al. 

have formulated PLGA nanoparticles loaded with indocyanine green (ICG) and determined their 

biodistribution in mice. The results demonstrated that the nanoparticle formulation significantly 

increased the ICG concentration and circulation time in plasma as well as the ICG uptake, accumulation 

and retention in various organs as compared to ICG solution. 54 Such a formulation can be 

explored in tumor-diagnosis as well as for PDT. 


It is essential to understand the molecular mechanisms involved in nanoparticle-mediated drug or 

gene delivery to explore their therapeutic potentials for cancer therapy. Also, the important questions 

we need to address are: What are the barriers in targeted cancer drug therapy? What is the 

efficiency of drug targeting with carrier systems? Can it provide a therapeutic dose of the drug in the 

tumor tissue? How long should the drug effect in the tumor tissue be sustained to regress it 

completely? Is a combination of drugs, especially those that work by different pathways, more 

effective than single-drug therapy? These and other questions are critical as we move from a 

conceptual stage to reality. With better understanding of molecular targets, discovery of more 

potent drugs and simultaneous developments in nanotechnology it seems that an effective cancer 

therapy is forthcoming. 



Grant support from the National Institutes of Health (R01 EB003975) is gratefully acknowledged. 

The authors also thank Ms. Elaine Payne and Ms. DeAnna Loibl for providing 

administrative support. 



REFERENCES 

1. Greish, K. et al., Macromolecular therapeutics: Advantages and prospects with special emphasis on 

solid tumour targeting, Clin. Pharmacokinet., 42, 1089, 2003. 

2. Luo, Y. and Prestwich, G. D., Cancer-targeted polymeric drugs, Curr. Cancer Drug Targets, 2, 209, 

2002. 

3. Fang, J., Sawa, T., and Maeda, H., Factors and mechanism of “EPR” effect and the enhanced antitumor 

effects of macromolecular drugs including SMANCS, Adv. Exp. Med. Biol., 519, 29, 2003. 

4. Sahoo, S. K. and Labhasetwar, V., Enhanced antiproliferative activity of transferrin-conjugated 

paclitaxel-loaded nanoparticles is mediated via sustained intracellular drug retention, Mol. Pharm., 

2, 373, 2005. 

5. Saeki, T. et al., Drug resistance in chemotherapy for breast cancer, Cancer Chemother. Pharmacol., 

56(Suppl 7), 84, 2005. 

6. Kabanov, A. V., Batrakova, E. V., and Miller, D. W., Pluronic block copolymers as modulators of 

drug efflux transporter activity in the blood–brain barrier, Adv. Drug Deliv. Rev., 55, 151, 2003. 

7. Kabanov, A. V., Batrakova, E. V., and Alakhov, V. Y., Pluronic block copolymers for overcoming 

drug resistance in cancer, Adv. Drug Deliv. Rev., 54, 759, 2002. 

8. Jain, K. K., Nanotechnology-based drug delivery for cancer, Technol. Cancer Res. Treat., 4, 407, 

2005. 

9. Singh, K. K., Nanotechnology in cancer detection and treatment, Technol. Cancer Res. Treat., 4, 583, 

2005. 

10. Meisheid, A. M., Targeted therapies in the treatment of cancer, J. Contin. Educ. Nurs., 36, 193, 2005. 

11. Reddy, L. H., Drug delivery to tumours: Recent strategies, J. Pharm. Pharmacol., 57, 1231, 2005. 

12. Adams, J. D. et al., Taxol: a history of pharmaceutical development and current pharmaceutical 

concerns, J. Natl. Cancer Inst. Monogr., 15, 141, 1993. 

13. Sahoo, S. K. and Labhasetwar, V., Nanotech approaches to drug delivery and imaging, Drug Discov. 

Today, 8, 1112, 2003. 

14. Maeda, H. et al., Tumor vascular permeability and the EPR effect in macromolecular therapeutics: 

a review, J. Control. Release, 65, 271, 2000. 

15. Maeda, H., Sawa, T., and Konno, T., Mechanism of tumor-targeted delivery of macromolecular drugs, 

including the EPR effect in solid tumor and clinical overview of the prototype polymeric drug 

SMANCS, J. Control. Release, 74, 47, 2001. 

16. Folkman, J. and Shing, Y., Angiogenesis, J. Biol. Chem., 267, 10931, 1992. 

17. Jain, R. K., Integrative pathophysiology of solid tumors: Role in detection and treatment, Cancer 

J. Sci. Am., 4(Suppl 1), S48, 1998. 

18. Jain, R. K., Delivery of molecular and cellular medicine to solid tumors, J. Control. Release, 53, 49, 

1998. 


tissue, Adv. Drug Deliv. Rev., 55, 329, 2003. 

20. Minko, T. et al., Molecular targeting of drug delivery systems to cancer, Curr. Drug Targets, 5, 389, 

2004. 

21. Gref, R. et al., Biodegradable long-circulating polymeric nanospheres, Science., 263, 1600, 1994. 

22. Gref, R. et al., Poly(ethylene glycol)-coated nanospheres: Potential carriers for intravenous drug 

administration, Pharm. Biotechnol., 10, 167, 1997. 

23. Vasir, J. K. and Labhasetwar, V., Targeted drug delivery in cancer therapy, Technol. Cancer Res. 

Treat., 4, 363, 2005. 

24. Sahoo, S. K., Ma, W., and Labhasetwar, V., Efficacy of transferrin-conjugated paclitaxel-loaded 

nanoparticles in a murine model of prostate cancer, Int. J. Cancer, 112, 335, 2004. 

25. Jain, R. A., The manufacturing techniques of various drug loaded biodegradable poly(lactide-coglycolide) 

(PLGA) devices, Biomaterials, 21, 2475, 2000. 

26. Shive, M. S. and Anderson, J. M., Biodegradation and biocompatibility of PLA and PLGA microspheres, 

Adv. Drug Deliv. Rev., 28, 5, 1997. 

27. Hanafusa, S. et al., Biodegradable plate fixation of rabbit femoral shaft osteotomies. A comparative 

study, Clin. Orthop. Relat. Res., 262, 1995. 


250 


28. Matsusue, Y. et al., Tissue reaction of bioabsorbable ultra high strength poly (L-lactide) rod. A longterm 

study in rabbits, Clin. Orthop. Relat. Res., 246, 1995. 

29. Langer, R., Tissue engineering: a new field and its challenges, Pharm. Res., 14, 840, 1997. 

30. Mooney, D. J. et al., Long-term engraftment of hepatocytes transplanted on biodegradable polymer 

sponges, J. Biomed. Mater. Res., 37, 413, 1997. 

31. Guzman, L. A. et al., Local intraluminal infusion of biodegradable polymeric nanoparticles. A novel 

approach for prolonged drug delivery after balloon angioplasty, Circulation, 94, 1441, 1996. 

32. Kim, S. H. et al., Target-specific cellular uptake of PLGA nanoparticles coated with poly(L-lysine)– 

poly(ethylene glycol)–folate conjugate, Langmuir, 21, 8852, 2005. 

33. Mo, Y. and Lim, L. Y., Paclitaxel-loaded PLGA nanoparticles: Potentiation of anticancer activity by 

surface conjugation with wheat germ agglutinin, J. Control. Release, 108, 244, 2005. 

34. Cegnar, M. et al., Poly(lactide-co-glycolide) nanoparticles as a carrier system for delivering cysteine 

protease inhibitor cystatin into tumor cells, Exp. Cell Res., 301, 223, 2004. 

35. Sanchez, A. et al., Biodegradable micro- and nanoparticles as long-term delivery vehicles for 

interferon-alpha, Eur. J. Pharm. Sci., 18, 221, 2003. 

36. Yoo, H. S. et al., In vitro and in vivo anti-tumor activities of nanoparticles based on doxorubicin– 

PLGA conjugates, J. Control. Release, 68, 419, 2000. 

37. Yamamoto, M. and Curiel, D. T., Cancer gene therapy, Technol. Cancer Res. Treat., 4, 315, 2005. 

38. Seth, P., Vector-mediated cancer gene therapy: an overview, Cancer Biol. Ther., 4, 512, 2005. 

39. Maitland, N. J., Stanbridge, L. J., and Dussupt, V., Targeting gene therapy for prostate cancer, Curr. 

Pharm. Des., 10, 531, 2004. 

40. Vasir, J. K. and Labhasetwar, V., Polymeric nanoparticles for gene delivery, Expt. Opin. Drug Del., 

3, 325, 2006. 

41. Panyam, J. et al., Rapid endo-lysosomal escape of poly(DL-lactide-co-glycolide) nanoparticles: 

Implications for drug and gene delivery, FASEB J., 16, 1217, 2002. 

42. Prabha, S. et al., Size-dependency of nanoparticle-mediated gene transfection: Studies with fractionated 

nanoparticles, Int. J. Pharm., 244, 105, 2002. 

43. Prabha, S. and Labhasetwar, V., Critical determinants in PLGA/PLA nanoparticle-mediated gene 

expression, Pharm. Res., 21, 354, 2004. 

44. Csaba, N. et al., PLGA: Poloxamer and PLGA:Poloxamine blend nanoparticles: New carriers for gene 

delivery, Biomacromolecules, 6, 271, 2005. 

45. Prabha, S. and Labhasetwar, V., Nanoparticle-mediated wild-type p53 gene delivery results in 

sustained antiproliferative activity in breast cancer cells, Mol. Pharm., 1, 211, 2004. 

46. He, Q. et al., Preparation and characteristics of DNA-nanoparticles targeting to hepatocarcinoma 

cells, World J. Gastroenterol., 10, 660, 2004. 

47. Kumar, M. N. et al., Cationic poly(lactide-co-glycolide) nanoparticles as efficient in vivo gene transfection 

agents, J Nanosci. Nanotechnol., 4, 990, 2004. 

48. Dolmans, D. E., Fukumura, D., and Jain, R. K., Photodynamic therapy for cancer, Nat. Rev. Cancer,3, 

380, 2003. 

49. Huang, Z., A review of progress in clinical photodynamic therapy, Technol. Cancer Res. Treat., 4, 

283, 2005. 

50. Konan, Y. N. et al., Encapsulation of p-THPP into nanoparticles: Cellular uptake, subcellular 

localization and effect of serum on photodynamic activity, Photochem. Photobiol., 77, 638, 2003. 

51. Vargas, A. et al., Improved photodynamic activity of porphyrin loaded into nanoparticles: an in vivo 

evaluation using chick embryos, Int. J. Pharm., 286, 131, 2004. 

52. Konan, Y. N. et al., Enhanced photodynamic activity of meso-tetra(4-hydroxyphenyl)porphyrin by 

incorporation into sub-200 nm nanoparticles, Eur. J. Pharm. Sci., 18, 241, 2003. 

53. Konan, Y. N. et al., Preparation and characterization of sterile sub-200 nm meso-tetra (4-hydroxylphenyl)porphyrin-loaded 

nanoparticles for photodynamic therapy, Eur. J. Pharm. Biopharm., 55, 115, 

2003. 

54. Saxena, V., Sadoqi, M., and Shao, J., Polymeric nanoparticulate delivery system for Indocyanine 

green: Biodistribution in healthy mice, Int. J. Pharm., 308, 200, 2006. 


15 

CONTENTS 

Poly(Alkyl Cyanoacrylate) 

Nanoparticles for Delivery 

of Anti-Cancer Drugs 

R. S. R. Murthy and L. Harivardhan Reddy 

15.1 Cancer Therapy ................................................................................................................. 252 

15.2 Nanoparticles as Drug Delivery Systems ......................................................................... 252 

15.2.1 Polymeric Nanoparticles ..................................................................................... 253 

15.2.1.1 Nanoparticles Prepared by the Polymerization Process .................... 253 

15.2.1.1.1 Dispersion Polymerization .............................................. 253 

15.2.1.1.1 Emulsion Polymerization ................................................ 254 

15.2.1.2 Poly(Alkyl Cyanoacrylate) Nanoparticles ......................................... 254 

15.2.1.3 Poly(Alkyl Cyanoacrylate) Nanocapsules ......................................... 256 

15.2.2 Loading of Drugs to Poly(Alkyl Cyanoacrylate) Nanoparticles........................ 256 

15.2.3 Characterization of Nanoparticles....................................................................... 258 

15.2.3.1 Physicochemical Characterization ..................................................... 258 

15.2.3.2 Degradation of Poly(Alkyl Cyanoacrylate) Nanoparticles................ 260 

15.2.3.3 Influence of Enzymes on the Stability 

of Poly(Alkyl Cyanoacrylate) Nanoparticles..................................... 260 

15.2.3.4 Drug Release from Poly(Alkyl Cyanoacrylate) Nanoparticles ......... 261 

15.2.4 In Vivo Distribution of Poly(Alkyl Cyanoacrylate) Nanoparticles ................... 261 

15.2.4.1 Surface Modification to Alter Biodistribution................................... 264 

15.2.5 Toxicity of Polyalkyl Cyanoacrylate Nanoparticles........................................... 264 

15.2.5.1 Toxicity of Drugs Associated with Poly(Alkyl Cyanoacrylate) 

Nanoparticles ...................................................................................... 266 

15.2.6 Drug Delivery in Cancer with Poly(Alkyl Cyanoacrylate) Nanoparticles ........ 266 

15.2.6.1 Targeting to Cancer Cells and Tissues .............................................. 266 

15.2.6.2 Drug Delivery to Resistant Tumor Cells ........................................... 268 

15.2.6.3 Drug Delivery to Hepatocellular Carcinomas ................................... 272 

15.2.6.4 Drug Delivery to Brain Tumors......................................................... 274 

15.2.6.5 Drug Delivery to Breast Cancer ........................................................ 275 

15.2.6.6 Drug Delivery to Lymphatic Carcinomas ......................................... 276 

15.2.6.7 Drug Delivery to Gastric Carcinomas ............................................... 276 

15.2.6.8 Chemoembolization Using Alkylcyanoacrylates............................... 277 

15.2.6.9 Delivery of Peptide Anti-Cancer Drugs ............................................ 278 

15.3 Conclusions ....................................................................................................................... 280 

References .................................................................................................................................... 280 


251

252 

15.1 CANCER THERAPY 

Cancer is defined by two characteristics: uncontrolled cell growth and the ability to invade, metastasize, 

and spread to distant sites. Despite several modes of therapy—such as chemotherapy, 

immunotherapy, and radiotherapy—the therapy of cancer remains a challenge. Systemic 

chemotherapy is the widely adopted treatment for the disseminated malignant disease, while 

recent progress in drug therapy has resulted in curative chemotherapeutic regimens for several 

tumors. However, chronic administration of anti-cancer drugs leads to severe systemic toxicity. To 

improve the therapy with these anti-cancer agents and reduce this associated toxicity, drug delivery 

systems have been introduced to deliver the drugs directly to the site of interest. 1 Normal cells are 

also susceptible to the cytotoxic effects of chemotherapeutic agents and exhibit a dose response 

effect, but the response curve is shifted relative to that of malignant cells. Proliferative normal 

tissues such as bone marrow and gastrointestinal mucosa are generally the most susceptible 

to chemotherapy-induced toxicity. Because chemotherapy has been widely adopted for the treatment 

of cancer, several chemotherapeutic agents with potential anti-cancer activity have 

been introduced. 

Several drug delivery systems facilitate effective chemotherapy with the anti-cancer agents. 

These systems include liposomes, microparticles, supramolecular biovectors, polymeric micelles, 

and nanoparticles. Although research on targeted therapeutic systems for cancer therapy has not 

significantly contributed to their commercialization for human application, the introduction into the 

market of products like doxorubicin long-circulating liposomes (Doxil) and Polifeprosan-20 with 

carmustine (BCNU, Gliadel Wafer) for cancer therapy has renewed interest in the field of targeted 

drug delivery to cancers. 2 

15.2 NANOPARTICLES AS DRUG DELIVERY SYSTEMS 


The challenge of modern drug therapy is the optimization of the pharmacological action of the 

drugs coupled with the reduction of their toxic effects in vivo. The prime objectives in the design of 

drug delivery systems are the controlled delivery of the drug to its site of action at a therapeutically 

optimal rate and dosage to avoid toxicity and improve the drug effectiveness and therapeutic index. 

Among the most promising systems to achieve this goal are the colloidal drug delivery systems, 

which include liposomes, niosomes, microemulsions, and nanoparticles. Nanoparticles are 

considered promising colloidal drug carriers, as they overcome the technological limitations and 

stability problems associated with liposomes, niosomes, and microemulsions. For instance, 

camptothecin-based drugs, because of their poor solubility and labile lactone ring, pose challenges 

for drug delivery. Williams et al. 3 developed a nanoparticle delivery system for a camptotheca 

alkaloid (SN-38) that is stable in human serum albumin; high lactone concentrations were observed 

even after 3 h and showed prolonged in vivo half-life of the active (lactone) form. 

Nanoparticles are solid colloidal particles ranging in size from 10 to 1000 nm. Depending on 

the process used for their preparation, two types of nanoparticles are obtained: nanospheres and 

nanocapsules (Figure 15.1). Nanospheres possess a rigid matrix structure that incorporates the drug, 

whereas the nanocapsule contains an oily core incorporating the drug surrounded by a membrane 

structure. 4 These are prepared from either synthetic or natural macromolecules in which the drug is 

dissolved, entrapped, encapsulated, or adsorbed on the surface. Generally, the definition of nanoparticles 

includes not only the particles described by Birrenbach and Speiser 5 by the term 

“Nanopellets,” but also nanocapsules and polymer lattices such as the “molecular scale drug 

entrapment” products described by Rhodes et al. 6 and Boylan and Banker. 7 Nanoparticles 

possess a matrix structure that incorporates the pharmacologically active substance and facilitates 

the controlled release of the active agent. 


Poly(Alkyl Cyanoacrylate) Nanoparticles for Delivery of Anti-Cancer Drugs 253 

Polymeric matrix 

containing drug 

Nanosphere 

15.2.1 POLYMERIC NANOPARTICLES 

In recent years, biodegradable polymeric nanoparticles have attracted attention as potential drug 

carriers because of their applications in the controlled delivery of drugs, their ability to target 

organs and tissues, their function as carriers for DNA in gene therapy, and their ability to perorally 

deliver proteins, peptides, and genes. 8,9 Langer 8 stated that the construction of fully biomimetic 

matrices containing all of the required chemical and topographical factors has not yet been 

achieved. Scientists are focusing on the design of new polymeric materials in which both topographical 

features and chemical cues act synergistically to qualify as drug-delivery carriers. 

Polymeric nanoparticles can be prepared either by the polymerization of monomers or from the 

preformed polymers. This chapter will focus on the former process. 

15.2.1.1 Nanoparticles Prepared by the Polymerization Process 

Historically, the first methods used to produce nanoparticles were derived from the field of latex 

engineering developed by polymer chemists. These methods were based on the in situ polymerization 

of monomers in various media. In the early 1970s, Birrenbach and Speiser, the 

pioneers in the field, produced the first polymerized nanoparticles for pharmaceutical use. Two 

types of polymerization processes were adopted for the preparation of polymeric nanoparticles: 

dispersion polymerization (DP) and emulsion polymerization (EP). 

15.2.1.1.1 Dispersion Polymerization 

Oily core 

+ Drug 

Nanocapsule 

FIGURE 15.1 Representation of nanosphere and nanocapsule. 

Polymeric membrane 

A DP involves a monomer, an initiator, a solvent in which the newly formed polymer is insoluble, 

and a polymeric stabilizer. The polymer is formed in the continuous phase and precipitates into a 

new particle phase stabilized by the polymeric stabilizer. The nucleation is directly induced in the 

aqueous monomer solution. Small particles are formed by the aggregation of growing polymer 

chains precipitating from the continuous phase when these chains exceed a critical chain length. 

When enough stabilizer covers the particles, coalescence of these precursor particles with themselves 

and with their aggregates results in the formation of stable colloidal particles. In this 

technique, the monomer is dispersed or dissolved in an aqueous medium from which the formed 

polymer precipitates. 

Essentially, the initiation of a monomer for polymerization requires an initiator that can 

generate ions or radicals to start the polymerization process. If the nucleation of the monomer is 

due to ions, then the mechanism is called “ionic polymerization.” Depending on the type of ions 

produced, the ionic polymerization may be anionic or cationic. If the radical nucleates the 

monomer, then the mechanism is known as “radical polymerization.” 10 


254 

15.2.1.1.2 Emulsion Polymerization 

Emulsion polymerization is a method for producing a latex and a polymer that exhibit any desired 

morphology, composition, sequence distribution, surface groups, molecular weight distribution, 

particle concentration, or other characteristics. In this technique, the monomer is emulsified in a 

nonsolvent-containing surfactant, which leads to the formation of monomer-swollen micelles and 

stabilized monomer droplets. The polymerization is performed in the presence of an initiator. The 

initiator generates either radicals or ions depending on the type of initiator, and these radicals or 

ions nucleate the monomeric units and start the polymerization process. The monomer-swollen 

micelles exhibit sizes in the nanometer range and thus have a much larger surface area than the 

monomer droplets. It has been assumed that, once formed, the free reactive monomers would more 

probably initiate the reaction within the micelles itself. In this case, the monomer droplets would act 

as monomer reservoirs. Being slightly soluble in the surrounding phase, the monomer molecules 

reach the micelles by diffusion from the monomer droplets, thus allowing the polymerization to be 

continued in the micelles. 11,12 Generally, the reaction continues until the monomer completely 

disappears. The particles obtained by EP are usually smaller (100–300 nm) than the original 

stabilized monomer droplets in the continuous phase. EP may be performed using either organic 

or aqueous media as the continuous phase. 

Emulsion Polymerization in an Organic Continuous Phase. EP in the organic continuous phase 

was the first process reported for the preparation of nanoparticles. 5,13 In this process, the watersoluble 

monomers are polymerized. Acrylamide and the crosslinker N,N 0 -bisacrylamide were the 

prototype monomers to be polymerized using chemical initiators such as N,N,N 0 ,N 0 -tetramethylethylenediamine 

and potassium peroxdisulphate or by light irradiation by g- or UV-radiations. 

However, the high toxicity of monomers and the need for high amounts of organic solvents and 

surfactants limits the importance of this technique. 

The cyanoacrylate monomers are relatively less toxic than acrylamide, and their polymers are 

biodegradable. Poly(alkyl cyanoacrylate) (PACA) nanoparticles were prepared by EP in the 

continuous organic phase. In this case, the monomer was added to the organic phase, creating 

nanoparticles with a shell-like structure (nanocapsules), as well as solid particles. 14 The mechanism 

of this nanocapsule formation was explained as following. The drug dissolved in aqueous phase was 

solubilized in the organic phase (containing surfactants) such as iso-octane, cyclohexane–chloroform, 

isopropyl myristate–butanol, and hexane. This result in a microemulsion with water-swollen 

micelles containing the drug. The monomer diffuse into the swollen micelles, and the OH K ions 

initiated the polymerization. The polymerization is so rapid that only an impermeable wall could be 

formed at the organic/water interface, preventing the diffusion of further monomers into the 

particles. The high amounts of organic solvents and surfactants required for this process, 

however, has greatly limited its application. 

Emulsion Polymerization in an Aqueous Continuous Phase. This technique is widely used for 

the preparation of nanoparticles by the polymerization of a wide variety of monomers, including 

alkylcyanoacrylates. Employing very low quantities of surfactants, it is only used to stabilize the 

newly formed polymer particles. Apart from this, the polymerization of alkylcyanoacrylates has 

also been carried out in the absence of surfactants, 15 in aqueous media containing dextran or 

hydroxylpropyl-b-cyclodextrin (HPCD). Poly(ethyl cyanoacrylate) and poly(isobutyl cyanoacrylate) 

(PIBCA) nanospheres containing metaclopramide have been prepared using this technique. 

However, the resulting drug loading was only 9.2% and 14.8%, respectively. 

15.2.1.2 Poly(Alkyl Cyanoacrylate) Nanoparticles 


The first report on the synthesis of PACA nanoparticles appeared in 1979. 16 PACA nanoparticles 

are biodegradable and hence are eliminated rapidly from the body. 17 The cyanoacrylate monomers 

are polymerized in the aqueous medium by the anionic polymerization; 16,18 the mechanism is 



OH− 

+ 

H2C 

N 

C 

C 

C 

O 

O 

R1 

HO 

H2C N 

C 

C 

C 

O 

O 

R1 

HO 

H2C 

depicted in Figure 15.2. The cyanoacrylate monomer is initiated by OH K ions, and the polymerization 

was performed at a low pH to control the reaction, which otherwise would have led 

to rapid polymerization resulting in the precipitation of large polymer aggregates. The stabilizers 

used were found to have a significant influence on the size of the formed polymer particles. High 

concentrations of poloxamer 188 (above 2%) reduced the particle size of polyisobutyl cyanoacrylate 

nanoparticles from around 200 nm without an emulsifier, to as low as 31–56 nm. 19 

Nanoparticles prepared without stabilizers or prepared using polysorbates were found to have 

monomodal molecular weight distribution, with mean molecular weights of 1000–4000 Da, 

whereas the stabilizers such as dextrans or poloxamers led to a distinctive bimodal distribution 

of the molecular weights, with peaks at 1000–4000 and 20,000–40,000 Da. This bimodal distribution 

is indicative of two separate polymerization reactions. The first reaction probably occurs in 

the aqueous phase where the termination of polymerization by H C ions is rapid, leading to small 

molecular weights. The second reaction is possibly the polymerization of captured growing unterminated 

polymer molecules within the primary particles. Due to the low concentration of H C ions 

in this environment, the termination frequency is reduced and hence the molecular weights would 

be much higher. Due to the initiation of cyanoacrylate polymerization by bases, the basic drugs also 

can act as initiators. However, because of the complexity of cyanoacrylate polymerization, the 

resulting particle size in such multi-component systems is difficult to predict. 

Poly(butyl cyanoacrylate) (PBCA) nanoparticles of n-butyl cyanoacrylate containing methotrexate 

were prepared by DP and EP. 20 DP nanoparticles were prepared using dextran as a 

stabilizer; the EP nanoparticles were stabilized by poloxamer 188. A high zeta potential was 

observed for nanoparticles prepared by the DP method, whereas the incorporation of methotrexate 

resulted in a decrease in zeta potential. The DP nanoparticles exhibited a high release of methotrexate, 

suggesting the channelizing effect of dextran chains incorporated into nanoparticles during 

polymerization. When tested in two different release media such as 0.1 mol L K1 HCl and pH 7.4 

phosphate buffer, a significant difference (p!0.01) in release rates was found for DP and EP 

nanoparticles. Drug release from both the nanoparticles followed Fickian diffusion in 0.1mol 

L K1 HCl, whereas the mechanism was found anomalous in pH 7.4 phosphate buffer. 

A similar study conducted on PBCA nanoparticles loaded with doxorubicin hydrochloride by 

incorporation and adsorption techniques 21 showed rapid drug release in 0.001-N HCl from both 

DP and EP nanoparticles; the release kinetics followed the Higuchi equation. 

N 

C 

C− 

C 

O 

O 

R1 

H+ Monomer 

n H 

FIGURE 15.2 Mechanism of alkyl cyanoacrylate polymerization. (From Behan, N., Birkinshaw, C., and 

Clarke, N., Biomaterials, 22, 1335, 2001. With permission.) 


256 

15.2.1.3 Poly(Alkyl Cyanoacrylate) Nanocapsules 

Nanocapsules are spherical structures formed by an envelope surrounding a liquid central cavity. 

The technique for the preparation of PIBCA nanocapsules was first reported by Fallaouh et al. 22 

The nanocapsules were obtained by injecting the organic phase, consisting of oil, with the dissolved 

drug, isobutylcyanoacrylate, and ethanol into the aqueous phase containing Pluronic F68 (a 

nonionic surfactant) under magnetic stirring. The colloidal suspension was concentrated by evaporation 

under a vacuum and filtered to obtain nanocapsules that possessed a shell-like structure 

with a wall 3-nm thick. This thickness was confirmed by transmission electron microscopy (TEM) 

after negative staining with phosphotungstic acid. The pH of the polymerization medium played an 

important role in obtaining nanocapsules. Nanocapsules with uniform size and low polydispersity 

were obtained only between pH 4 and 10. The concentration of alcohol above 15% only led to the 

formation of isolated nanocapsules. The mechanism of nanocapsule formation is probably the 

interfacial polymerization described by Florence et al. 23 

for the manufacture of 

PACA microcapsules. 

A different method was adopted by Chouinard et al. 24 for the preparation of poly(isohexyl 

cyanoacrylate) (PIHCA) nanocapsules. The monomer was treated with sulfur dioxide and dissolved 

in a solution containing Miglyol 810 (caprylic/capric triglyceride) in ethanol. This solution was 

added slowly, in a 1:2 ratio, into an aqueous phase containing a solution of 0.5% poloxamer 407 and 

10 mM phosphate buffer under stirring. The nanocapsules were then purified by centrifugation and 

washing with distilled water. In this method, the sulfur dioxide acts as an inhibitor, preventing the 

polymerization of the cyanoacrylates by the ethanol. With the addition of the oily phase to 

the aqueous phase, the sulfur dioxide and ethanol migrate to the external phase, resulting in the 

formation of a very fine emulsion. In this case, the nanocapsule size was mainly influenced by 

Miglyol concentration. Freeze-fracture electron microscopy studies showed that polymeric walls of 

different thicknesses and densities can be prepared, which could be significant for the design of 

nanocapsules with tailored drug-release kinetics. 

Palumbo et al. 25 adopted a similar technique for the preparation of polyethyl-2-cyanoacrylate 

nanocapsules containing idebenone (an antioxidant). In this method, the acetonic solution of 

Miglyol 812, idebenone, and monomer was added to 100 mL of aqueous phase (pH 7) containing 

Tween 80. The presence of nonionic surfactant allowed the polymerization of the ethylcyanoacrylate 

at the oil–water interface, thus encapsulating Miglyol 812 droplets. The immediate 

polymerization triggered the formation of drug-loaded nanocapsules, with the suspensions concentrated 

under a vacuum to remove any acetone. Tween 80 reduced the hydrodynamic size of the 

emulsion droplets as a function of its concentration, resulting in smaller-sized nanocapsules. These 

nanocapsules exhibited a negative charge that was determined by zeta-potential measurements. The 

freeze-fracture electron microscopy revealed the presence of internal oil droplets surrounded by the 

polymeric shell of the nanocapsule. 

15.2.2 LOADING OF DRUGS TO POLY(ALKYL CYANOACRYLATE) NANOPARTICLES 


The drugs were loaded to PACA nanoparticles, either by incorporation during the polymerization 

process or by absorption onto the surface of the preformed particles. The former case can lead to the 

covalent coupling of the drug to the polymer, 26 or to the formation of a solid solution or solid 

dispersion of the drug in the polymer network. The addition of the drug to empty nanoparticles may 

lead to covalent coupling 27 of the drug; alternatively, the drug may be bound by sorption. The 

sorption can also lead either to the diffusion of the drug into the polymer network and the formation 

of a solid solution, 28 or to surface adsorption of the drug. 29 In the majority of cases, the type of drug 

loading will determine the carrier capacity. The carrier capacity is defined as the percentage of drug 

associated with a given amount of nanoparticles for a given initial concentration of drug without 

considering the characteristics of the adsorption isotherm. 30,31 Several drugs such as betaxolol 



chlorhydrate, 32 hematoporphyrin, 33 and primaquine 34 were adsorbed onto the preformed PACA 

nanoparticles, whereas drugs such as doxorubicin 35 and actinomycin D were incorporated into the 

nanoparticles during the polymerization process. The incorporation of rose bengal into the poly 

(butyl-2-cyanoacrylate) nanoparticles during the polymerization led to much greater carrier 

capacity than did the simple adsorption method. 29 The rose bengal incorporated into nanoparticles 

during polymerization is expected to be either dissolved, dispersed in, or adsorbed on the polymer 

matrix and remaining to be adsorbed on the particle surface. In contrast, after the adsorption 

process, the drug solely occupies the external surface of the nanoparticles, including minor pores 

and pits. Adsorption of hematoporphyrin onto PIBCA nanoparticles resulted in poor carrier 

capacity of the nanoparticles, because the drug was adsorbed mainly at the surface of the nanoparticles. 

33 A similar report by Beck et al. 28 describes the preparation of PBCA nanoparticles 

loaded with mitoxantrone using both the incorporation and adsorption methods. The proportion 

of mitoxantrone bound to the particles was analyzed to be about 15% of the initial drug concentration 

with the incorporation method and about 8% with the adsorption method. 

The type of drug binding onto the nanoparticles may result in different release mechanisms and 

rates. 36 Whether the drug is present in the form of a solid solution or a solid dispersion, its release 

characteristics mainly depend on the degradation rate of the polymer. 30 Drug desorption (tested in 

phosphate buffered medium pH 5.9) resulted essentially from a shift in the equilibrium conditions 

between free and adsorbed drug following the dilution of the nanoparticle suspension, modification 

of the pH conditions, or enzymatic hydrolysis by the added esterases. The release of hematoporphyrin 

was very rapid at physiological temperature and pH, and hence would not be useful for 

in vivo applications. 33 Finally, the surface charge of the particles and binding type between the drug 

and nanoparticles are the important parameters that determine the rate of desorption of the drug in 

the nanoparticles. 32 Recently, Poupaert and Couvreur 37 developed a computational derived structural 

model of doxorubicin interacting with oligomeric PACA in nanoparticles. The oligomeric 

PACAs are highly lipophilic entities and scavenge the amphiphilic doxorubicin during the polymerization 

process by extraction of protonated species from the aqueous environment to an 

increasingly lipophilic phase embodied by the growing PACAs. Interesting, hydrogen bonds 

were established between the N and H function of doxorubicin and the cyano groups of alkylcyanoacrylate. 

Therefore, the cohesion in this assembly comes from a blend of dipole-charge 

interaction, H-bonds, and hydrophobic forces. Upon hydrolytic erosion of the nanoparticles involving 

mainly the hydrolysis of ester groups to carboxylate, the assembly of doxorubicin–PACA 

tends to loosen the cohesion as the hydrophobic forces decline due to water intrusion and the 

increasing contribution of repulsive forces between anionic carboxylates. This creates, for 

example, a high local gradient of doxorubicin at the cell membrane surface, permitting MDR 

efflux to be overwhelmed. Very recently, heparin–PIBCA copolymer nanoparticles were developed 

as a carrier for hemoglobin, 38 which was indicated for treatment in thrombosis oxygen-deprived 

pathologies. In water, these copolymers spontaneously form nanoparticles with a ciliated surface of 

heparin. The heparin bound to the nanoparticle preserved its antithrombotic activity. The bound 

hemoglobin also maintained its capacity to bind ligands. 

The integration of drugs with polymerization additives was observed in the case of 5-fluorouracil 

(5-FU)-loaded PBCA nanoparticles. 39 5-FU in acidic solution (pH 2–3) may interfere in the 

initiation process through its amino groups via the formation of zwitterions. The proposed zwitterionic 

mechanism of initiation was supported by the molecular weight profiles of the polymer, as 

determined by gel permeation chromatography, and the covalent linkage of the cytostatic to the 

main polymer chain. 1 H NMR analyses demonstrated that a significant fraction of 5-FU was 

covalently bonded to the PBCA chains through its amino groups, preferentially through one of 

the two nitrogen atoms. 

Stella et al. 40 developed a new concept to target the folate-binding protein by designing poly 

(ethylene glycol) (PEG)-coated biodegradable nanoparticles coupled to folic acid. This system is 

the soluble form of the folate receptor that is over expressed on the surface of many tumoral cells. 


258 

The copolymer poly[aminopoly(ethylene glycol)cyanoacrylate-co-hexadecyl cyanoacrylate] 

[poly(H 2NPEGCA-co-HDCA)] was synthesized, and their nanoparticles were prepared using 

the nanoprecipitation technique. The nanoparticles were then conjugated to the activated folic 

acid via PEG terminal amino groups and purified from unreacted products. The specific 

interaction between the conjugate folate-nanoparticles and the folate-binding protein was 

confirmed by surface plasmon resonance. This interaction did not occur with the nonconjugated 

nanoparticles used as a control. 

15.2.3 CHARACTERIZATION OF NANOPARTICLES 

15.2.3.1 Physicochemical Characterization 

The methods used for the physicochemical characterization of nanoparticles are listed in 

Table 15.1. Particle size is the prominent feature of nanoparticles; the fastest and most routine 

methods of size analysis are photon correlation spectroscopy (PCS) and laser diffractometry (LD). 

The former method is useful for the determination of smaller particles, 41 whereas the latter method 

is useful for the determination of larger particles. PCS determines the hydrodynamic diameter of the 

nanoparticles via Brownian motion. The electron microscopy methods also allow the exact particle 

determination. Scanning electron microscopy requires the coating of the dry sample with a conductive 

material such as gold. The gold coating usually results in an estimate of particle sizes that is 

slightly greater than the normal. TEM, with or without staining, is a relatively easier method to 

determine particle size. Some nanoparticle materials are not electron dense and thus cannot be 

stained; others melt and sinter when irradiated by the electron beam of the microscope and thus 

cannot be visualized by this method. Such nanoparticles can be visualized in TEM after freeze 

fracture or freeze substitution. 42 This method is optimal because it also allows for the fracturing 

of the particles and consequently an observation of their interior. Unfortunately, the method is 

time-consuming and cannot be used for the routine determinations. Recently, new types of highresolution 

microscopes such as the atomic-force microscope, the laser-force microscope, and the 

scanning-tunneling microscope were developed. 43–46 These microscopes are especially useful for 

the investigation of nanoparticle surfaces. 

The molecular weight determination of nanoparticles can be performed after the particles have 

been dissolved in an appropriate solvent and analyzed by gel-permeation chromatography. 

Accurate results can be obtained only if the polymer standards have a similar structure and 

similar properties as the test material. The limitation of this technique, however, is that the 

TABLE 15.1 

List of Characterization Techniques for Nanoparticles 

Parameter Method of Characterization 


Particle size Photon correlation spectroscopy, laser diffraction, transmission electron microscopy, 

scanning electron microscopy 

Molecular weight Gel permeation chromatography 

Density Helium compression pycnometer 

Crystallinity X-ray diffractometry, differential scanning calorimetry 

Surface charge Electrophoresis, laser Doppler anemometry, amplitude-weighted phase structuration 

Hydrophobicity Hydrophobic interaction chromatography, contact angle measurements 

Surface characteristics Scanning electron microscopy, atomic force microscopy 

Surface element analysis X-ray photoelectron spectroscopy for chemical analysis (ESCA) 



molecular weight of cross-linked polymers and nanoparticles from natural macromolecules cannot 

be determined. 

Density measurements can be performed by the helium compression pycnometry and by 

density gradient centrifugation. A comparison of these two methods may offer information about 

the internal structure of nanoparticles. 

Further information about the nanoparticle structure and crystallinity may be obtained using 

x-ray diffraction and thermoanalytical methods such as differential scanning calorimetry, differential 

thermal analysis, thermogravimetry, and thermal optical analysis. 47 Thechargeon 

nanoparticle surfaces is mainly determined by electrophoretic mobility, laser Doppler anemometry, 

and amplitude-weighted phase structuration. 

Hydrophobicity of the nanoparticles can be determined by two major methods: contact angle 

measurements and hydrophobic interaction chromatography. The former can be performed only on 

flat surfaces and not directly on hydrated nanoparticles in their dispersion media. For this reason, 

hydrophobic interaction chromatography is the better method, wherein a differentiation between 

nanoparticles with different surface properties can be obtained by loading the particles on columns 

with alkyl-sepharose and eluting them with a Triton X-100 gradient. 48 

Kreuter 41 performed the physicochemical characterization of polymethyl, polyethyl, and 

PBCA nanoparticles. Because it gives only the mean diameter and the multimodal size distributions 

cannot be measured, determination of nanoparticle size by Photon Correlation Spectrometry is very 

susceptible to errors caused by the presence of bigger particles in the dispersion. Another disadvantage 

is that this method measures the Brownian motion of the particles, and the particle 

diameter is then calculated from the measurements via the diffusion coefficient. 41 Thus, the particle 

size determined by this method can be affected by the surrounding medium, including the adsorbed 

surfactants or hydration layers, and may therefore be different from the sizes obtained by electron 

microscopy. In case of characterization using electron microscopy, the individual particles can be 

analyzed and measured. According to Kopf et al., 26 the surfactants present after the separation of 

the dispersion medium from the nanoparticles will coat the nanoparticles and lead to the disappearance 

of individual nanoparticle structures. This coating significantly reduces the particle 

surface area, and polycyanoacrylate nanoparticles could thus not be obtained surfactant-free in a 

nonaggregated form. The nanoparticles were x-ray amorphous, and their diffractograms indicated 

no sign of crystallinity. 20,21 The surface charge of colloidal particles expressed by their electrophoreitc 

mobility may have an important influence on their body distribution behavior in humans 

and animals. 49,50 The PACA nanoparticles possess negative charge. Charge measurements in serum 

yielded a more pronounced decrease caused by the adsorption of the serum contents. Another 

parameter that is of great importance for the fate of nanoparticles in the blood is the hydrophilicity/hydrophobicity 

of the particle surface. The latter has a profound influence on the interaction of 

the particles with the blood components. 49,51,52 The particles’ hydrophilicity/hydrophobicity can be 

determined by measuring the water contact angle. 53 This was done by centrifugation, followed by 

washing with water and drying the poly(cyanoacrylate) nanoparticles. The dried material was 

dissolved in acetone and casted as a film with subsequent solvent evaporation. 41 The water 

droplet contact angles for polymethyl, polyethyl, and PBCA nanoparticles were reported as 

55.68, 64.78, and 68.98, respectively. 41 

Muller et al. 48 studied the physicochemical properties of alkyl cyanoacrylates of different chain 

lengths such as methyl, ethyl, isobutyl, and isohexyl cyanoacrylate. Zeta potential of the nanoparticles 

was determined as an alternative means of evaluating particle surface charge. Interaction of 

nanoparticles with serum proteins was assessed by the zeta-potential measurements. The nanoparticles 

were incubated with poloxamer 407 (Pluronic F-108), and the hydrophobicity determination 

by contact angle measurement revealed that the coating layer was very thin. The layer thickness 

was higher in isohexyl cyanoacrylate particles, indicating that they had the most hydrophobic 

nature of all the polymers used. 


260 

The molecular weight differences between polycyanoacrylate nanoparticles with different sidechain 

lengths as well as the molecular weight differences of polycyanoacrylate nanoparticles 

produced at different hydrochloric acid concentrations or surfactant concentrations appear to not 

be very pronounced. 54 Nevertheless, an increase in side-chain length led to the increasing molecular 

weights, whereas increasing the hydrochloric acid concentrations reduced the molecular weights. 

15.2.3.2 Degradation of Poly(Alkyl Cyanoacrylate) Nanoparticles 


The type of degradation (bulk or surface degradation) of polymethyl-, polyethyl-, and poly(isohexyl 

cyanoacrylate) nanoparticles was determined using photon correlation spectroscopic measurements. 

48 In the case of surface degradation, the particle size should show an immediate increase 

while the polydispersity index remaining constant. If bulk degradation dominates, a lag period 

should occur, preceding the decrease in mean size due to disintegration of the particles. This 

process of disintegration would lead to a more heterogeneous distribution of particle sizes and 

consequently to an increase in the polydispersity index. Incubation of polyethyl cyanoacrylate 

nanoparticles with in 10 K4 N NaOH led to an immediate, continuous decrease in particle size, 

as well as an unchanged polydispersity index indicating the predominant surface degradation. 

Incubation of PIHCA nanoparticles in 10 K4 N NaOH led to no detectable size decrease; polydispersity 

was also unchanged, indicating the degradation by surface erosion at a much slower rate 

than poly(ethyl cyanoacrylate) nanoparticles. Coating the particles with poloxamers did not accelerate 

the particle degradation. At low electrolyte concentration, the size and dispersity of PIHCA 

nanoparticles remained unchanged during incubation in 10 K4 N NaOH. At high ionic strength, the 

size increase of nanoparticles as a result of flocculation was larger than the size decrease due to 

degradation. In vitro degradation was also determined by turbidimetric measurements. 55 Poly(methyl 

cyanoacrylate) and poly(ethyl cyanoacrylate) nanoparticles underwent the fastest 

degradation. A high electrolyte concentration in the medium led to the formation of larger aggregates 

accompanied by an increase in absorption of dispersion. The slowly degrading polymers 

PIBCA and PIHCA probably release a low concentration of degradation products over a prolonged 

period of time, indicating their low toxicity. 

The stability studies on PBCA nanoparticles reveal that an acidic medium protects the nanoparticles 

against decomposition. 56,57 Higher temperatures promoted the degradation of 

nanoparticles. When incubated in human serum in vitro, the nanoparticles showed no particle 

agglomeration, suggesting their better tolerance in vivo. 

15.2.3.3 Influence of Enzymes on the Stability of Poly(Alkyl Cyanoacrylate) Nanoparticles 

The peroral route is the most convenient way of delivering drugs, leading to better patient compliance, 

especially during long-term treatment. However, many drugs, including proteins and 

peptides, are unstable in the gastrointestinal tract (GIT) or are insufficiently absorbed. Earlier 

studies have shown that cyanoacrylate nanoparticles 58,59 and nanocapsules 60,61 improve the absorption 

of several drugs like vincamin or insulin from the GIT. Maincent et al. 58 demonstrated the 

efficiency of PBCA nanoparticles as a drug delivery system for the GIT by improving the absorption 

of vincamin by more than 60% compared to an equimolar solution. Investigation of the 

parameters that influence the stability of nanoparticles in the GIT is much more difficult because 

of the complexity of the GIT. The pH changes from 2.0 in the stomach to 7–8 in the small intestine. 

Moreover, in different parts of the GIT there are different enzymes in differing concentrations. 

Degradation studies with cyanoacrylate was performed as early as 1972. 62 Lenaerts et al. 63 incubated 

PIBCA nanoparticles in rat liver microsomes. As these microsomes contain a mixture of 

enzymes, it was difficult to determine which enzyme took part in the degradation process. Scherer et 

al. 56 studied the influence of different enzymes in vitro on the stability of PBCA nanoparticles. The 

study used pepsin, amylase, and esterase enzymes, added in a buffer solution that contained 



dispersed nanoparticles. Amylase was selected because dextran is used as a stabilizer in nanoparticle 

preparation and was shown to be partially incorporated into the polymer chains. 64 Esterase was 

selected because hydrolysis of the ester side chain is the main route of degradation of PACA 

nanoparticles. 65 The study showed significant degradation of PBCA nanoparticles at 378C and 

was dependent on the pH of the buffer solution. The addition of pepsin and amylase to the 

buffer solution was found to have no influence on the stability of nanoparticles. The results upon 

addition of amylase revealed that dextran either was not incorporated to a significant degree or that 

its incorporation did not alter the regular nonenzymatic or esterase-dependant degradation pathway 

and velocity of this polymer in the form of nanoparticles. In contrast, the esterase significantly 

influenced the stability of the nanoparticles. Esterase demonstrated the highest activity of nanoparticle 

degradation at neutral pH values; as well, the degradation of PBCA nanoparticles was 

proportional to the amount of esterase added. These observations suggest that the degradation of 

PBCA nanoparticles does not occur before the particles reach the small intestine, as esterase is 

released by the pancreas into the lumen of the small intestine, the major site for absorption of 

several drugs. 

15.2.3.4 Drug Release from Poly(Alkyl Cyanoacrylate) Nanoparticles 

Various methods were used to study the in vitro drug release of nanoparticles. They include sideby-side 

diffusion cells with artificial or biological membranes, the dialysis-bag diffusion technique, 

the reverse dialysis sac technique, ultrafiltration, ultracentrifugation, and the centrifugal ultrafiltration 

techniques. Despite the development of these methods, some technical difficulties remain, 66 

including the separation of the nanoparticles from the release media when assessing the in vitro 

drug release from nanoparticles. Leavy and Benita 67 used a reverse dialysis method, in which the 

nanoparticles were added directly into the dissolution medium. The same technique was adopted by 

Calvo et al. 68 for the evaluation of nanoparticles, nanocapsules, and nanoemulsions. However, this 

method is not very sensitive for studying the rapid release formulations, although relatively 

speaking the dialysis bag technique is the best for assessing in vitro drug release. Because the 

nanoparticle dispersion is placed in the dialysis bag and immersed into the release medium, this 

technique avoids the problem of nanoparticle separation. 69,70 Considering the hindrance to drug 

diffusion caused by the dialysis bag, the release of the drug in solution form can be studied through 

the dialysis bag, and any diffusional barrier can be compensated for. 

15.2.4 IN VIVO DISTRIBUTION OF POLY(ALKYL CYANOACRYLATE) NANOPARTICLES 

Colloidal drug carriers such as liposomes and nanoparticles are able to modify the distribution of an 

associated substance. They can therefore be used to improve the therapeutic index of drugs by 

increasing their efficacy or reducing their toxicity. The polymer nanoparticles, particularly PACA 

nanoparticles, have attracted considerable attention as potential drug delivery systems, not only for 

their enhancement of therapeutic efficacy but also because they reduce the toxicity associated with a 

variety of drugs. Careful design of these delivery systems with respect to the target and route of 

administration may provide a solution to some of the delivery problems posed by many anti-cancer 

drugs, including doxorubicin and amphotericin B, 71 and new classes of active molecules such as 

peptides, proteins, genes, and oligonucleotides. 

The use of colloidal particulate carrier systems with controlled particle size (25 nm to 1.7 mmin 

diameter) is significant in such applications where diminished uptake by mononuclear phagocytes 

and specific targeting of carriers to particular tissues or cells is of great importance. Studies 

conducted on the tissue distribution of PBCA nanoparticles with a diameter of 127 nm, loaded 

with 1-(2-chloroethyl)-3-(1-oxyl-2,2,6,6-tetramethyl piperidinyl)-1-nitrosourea (spin-labeled nitrosourea, 

SLCNU) suspensions injected intraperitoneally (i.p.) into Lewis lung carcinoma-bearing 

mice showed a relatively low accumulation of nanoparticles in the liver and spleen; the 


262 


nanoparticles were mainly found in the lungs, kidneys, and heart. 72 The highest content of the 

particles was observed in the lungs of tumor-bearing experimental animals damaged by metastases, 

suggesting the possible usefulness of SLCNU–PBCA nanoparticles for targeting the 

lung metastases. 

Reddy and Murthy 73 compared the pharmacokinetics and tissue distribution of doxorubicin 

(Dox) solution with Dox-loaded PBCA nanoparticles synthesized by DP and EP techniques after 

intravenous (i.v.) and intraperitoneal (i.p.) injection in healthy rats. The elimination half-life (T1/2) 

and mean residence time (MRT) in blood was significantly higher than the free Dox after i.v. 

injection. After i.p. injection, DP nanoparticles quickly appeared in the blood and underwent rapid 

distribution to the organs of RES, whereas the EP nanoparticles were absorbed slowly into the 

blood and remained in the circulatory system for a longer time. The bioavailability (F) ofDP 

(w1.9-fold) and EP nanoparticles (w2.12-fold) was higher compared to that of the Dox solution 

after i.p. injection. The distribution to the heart of both types of nanoparticles was low after i.v. and 

i.p. injection, compared to Dox. This was further proved by experiments using nanoparticles 

radiolabeled with 99m Tc to study their distribution in Dalton’s lymphoma implanted in the thigh 

region of mice, when administered subcutaneously below the tumor region. 74 A significantly high 

tumor uptake of 99m Tc-EP nanoparticles (13-fold higher at 48 h post-injection) (P!0.001) was 

found compared to the 99m Tc-Dox solution. 

Blank [ 14 C]-poly(hexyl cyanoacrylate) nanoparticles with diameters of 200–300 nm injected 

intravenously into nude mice bearing a human osteosarcoma showed distribution to the liver, 

spleen, lung, heart, kidney, GI tract, gonads, brain, and muscle, as well as in serum and transplanted 

tumor fragments when investigated by liquid scintillation counting. 75 The peak levels in all organs 

with the exception of tumors and the spleen were reached within 24 h; the highest levels of 

radioactivity were found after 7 days in the tumor and spleen. The tissue distribution of naked 

and either normal immunoglobulin G or monoclonal antibody (antitumor osteogenic sarcoma)coated 

poly(hexyl-2-cyanoacrylate) nanoparticles in mice bearing human tumor xenografts showed 

similar results: the nanoparticles were deposited mainly in the liver and spleen, and no significant 

uptake was found in the tumors. 76 

The PACA nanoparticles exhibited rapid extravascular distribution after intravenous administration. 

Intravenous injection of isobutyl-2-cyanoacrylate nanoparticles resulted in high 

concentrations in the organs of the reticuloendothelial system such as the liver, spleen, and bone 

marrow. Similar observations were made by Kante et al. 77 in the case of PBCA nanoparticles 

injected intravenously in rats. In the case of mice with subcutaneously grafted Lewis lung carcinoma 

cells, the nanoparticles exhibited higher lung concentrations, indicating the use of 

nanoparticles in the treatment of lung metastasis. 17 However, the results were somewhat different 

for polybutyl-2-cyanoacrylate nanoparticles injected intravenously into rabbits. 78 The lung 

accumulation of nanoparticles was very low in that case, indicating that the nanoparticles 

(126 nm) were not mechanically filtered through the capillary bed of the lungs. The difference in 

lung concentrations compared to the observations of Grislain et al. 17 can be explained by the 

difference in particle size (254 nm). Coating with azidothymidine-loaded poly(hexyl cyanoacrylate) 

(PHCA) nanoparticles with polysorbate 80 led to higher brain levels of azidothymidine after 

intravenous injection in rats, indicating the facilitation of drug passage across the blood–brain 

barrier (BBB). 79 Physicochemical studies have shown that the coating of nanoparticles with 

block copolymers such as poloxamers and poloxamines induces a steric repulsion effect, minimizing 

the adhesion of particles to the surface of macrophages which in turn results in decreased 

phagocytic uptake and significantly higher levels in blood and non-RES organs, the brain, intestine, 

kidneys, etc. Borchard et al. 80 showed that coating the nanoparticles with surfactants resulted in 

uptake by brain endothelial cells, and that polysorbate 80 was the most effective surfactant for this 

purpose. Further investigations unequivocally demonstrated that polysorbate 80-coated PBCA 

nanoparticles loaded with enkaphalin analogue dalargin enabled the delivery of this drug across 

the BBB, achieving a significant pharmacological effect. Gulyaev et al. 81 reported significantly 



higher brain concentrations of intravenously injected polysorbate 80-coated Dox–PBCA nanoparticles 

in rats. Three probable mechanisms were proposed for the enhanced brain transport of 

polysorbate 80 coated nanoparticles: (1) endocytosis, (2) the opening of the tight endothelial 

junctions between the brain endothelial cells, 81 and (3) the inhibition of the P-glycoprotein 

efflux pump responsible for multi-drug resistance, which is a major obstacle to cancer 

chemotherapy. 82,83 

Isobutyl and isohexylcyanoacrylate nanoparticles were used by Ghanem et al. 84 as drug 

carriers, particularly for some anti-cancer drugs. Body distribution as well as pharmacokinetics 

in animals and partially in humans using a radiolabeled ( 111 In and 99m Tc) free drug and its nanoparticles 

showed 60–75% accumulation in the reticuloendothelial system. 

Skidan et al. 85 studied the antitumor efficacy of Dox loaded into polysorbate 80-coated nanoparticles 

to the brain. Pharmacokinetics were studied in rats after the intravenous injection of four 

Dox preparations: (1) solution in saline, (2) solution in polysorbate 80, intravenous (3) bound to 

nanoparticles, and (4) bound to nanoparticles coated with polysorbate 80. The results showed no 

significant difference in the body distribution between the two solution formulations, whereas the 

two nanoparticle formulations showed significantly decreased drug concentrations in the heart. 

High brain concentrations of Dox (O6 mg/g) were achieved with the nanoparticles overcoated 

with polysorbate 80 between 2 and 4 h, whereas the concentrations observed with the other three 

formulations were always negligible. Antitumor efficacy of Dox associated with polysorbate 

80-coated nanoparticles as assessed in rats with intracranially implanted 101/8 glioblastoma 

resulted in a significant 3.6-fold increase in survival, whereas administration of free Dox or 

blank nanoparticles did not modify the mortality when compared to the control group. 

Following the subcutaneous injections of metoclopramide solution (5 mg/kg) and three 

different metoclopramide nanosphere suspensions (10 mg/kg) on two phases in wistar rats, there 

was a rapid absorption, distribution, and elimination of the drug solution. The maximum drug 

concentration was observed after 30 min of subcutaneous injection of all the tested nanosphere 

formulations. 15 Polyethylcyanoacrylate-hydroxypropyl cyclodextrin showed the highest concentration, 

followed by polyisobutylcyanoacrylate–dextran and polyethylcyanoacrylate–dextran. The 

area under the curves of all nanoparticle formulations was 4.8- to 1.88-times higher than that 

obtained for the solution form. 

Earlier studies showed that polystyrene and other biodegradable microspheres delivered to the 

intestine are preferentially absorbed by the M-cells of the Peyer’s patches. 86,87 Intragastric administration 

of azidothymidine-loaded PIHCA nanoparticles resulted in high concentration in the 

intestinal mucosa as well as in the Peyer’s patches. 88 Once localized in the mucous or intestinal 

tissue, the azidothymidine could be slowly released by the enzymatic action of esterases. Very low 

levels of azidothymidine bound to nanoparticles were detected in the lymph, suggesting an efficient 

capture of particles through the Peyer’s patches resulting in an intracellular localization of the drug. 

As a result, some immune cells would carry azidothymidine to the mesenteric lymph nodes, the 

location of an important viral burden and a major replication site of HIV. 

PACA nanoparticles are proved to be suitable delivery systems in comparison to other colloidal 

systems, including liposomes. Reszka et al. 89 observed that the mitanxantrone-loaded PBCA nanoparticles 

showed the highest tumor concentrations in B-16 melanoma-bearing mice after intravenous 

injection. Four different formulations—mitaxantrone solution, mitaxantrone-loaded negatively 

charged liposomes (small unilamellar vesicles), 14 C-PBCA nanoparticles, and poloxamine 1508coated 

14 C-PBCA nanoparticles—were studied for the biodistribution and tumor accumulation in the 

tumor-bearing mice. The liposomal formulation showed prolonged circulation in the blood and 

higher accumulation in the liver and spleen after 24 h of injection, compared to the nanoparticle 

formulations. Despite the longer circulation in the blood, however, the liposomal formulation could 

not result in high concentrations in the tumors, compared to the nanoparticle formulations, which 

showed relatively lower circulation time in the blood, indicating the efficacy of PBCA nanoparticles 


264 

in delivering mitoxantrone to tumors. Similar results involving the higher efficacy of mitoxantroneloaded 

PBCA nanoparticles in B-16 melanoma was reported by Beck et al. 28 

15.2.4.1 Surface Modification to Alter Biodistribution 

The potential of injectable polymeric drug carriers (nanoparticles) is compromised by their 

accumulation in the tissues of the mononuclear phagocytic system (MPS). Thus, the therapeutic 

efficacy of the drugs associated with the nanoparticles is limited to the treatment of several liver 

diseases. 90,91 Long-circulating nanoparticles can be obtained by their surface modification with 

dysopsonic polymers such as PEG. 92,93 Indeed, these hydrophilic and flexible polymers can prevent 

the opsonin–nanoparticle interaction, which is the first step of the recognition by the immune 

system. In the tumor-bearing animal models, these long-circulating nanoparticles would be able 

to extravasate thorough the endothelium, allowing drugs to concentrate in the tumors. 94 Recently, it 

was shown that the synthesis of amphiphilic PEGylated polymers allows the direct preparation of 

PEG-coated nanoparticles, ensuring the stability of the PEG coating layer because the PEG chains 

remain chemically linked to the nanoparticle core. 95–97 Otherwise, a PEG coating layer simply 

adsorbed onto preformed nanoparticles is desorbed in vivo. 78 Peracchia et al. 98 synthesized a novel 

MePEGcyanoacrylate–hexadecylcyanoacrylate (PEG–PHDCA) copolymer, and the presence of a 

PEG layer on the nanoparticle surface was confirmed by surface chemical analysis. 99 Poly(alkyl 

cyanoacrylate) nanoparticles are rapidly biodegradable and hence could be able to release the 

associated drug into the bloodstream in a controlled manner before liver capture could occur. 

For solid tumors, extravasation could be followed by a rapid degradation of the polymer, leading 

to a rapid drug release. 100 The cytotoxicity of PHDCA polymer was decreased after its PEGylation 

in the mouse macrophage cell line J774. 101 This is due to the lower extent of the particle/cell 

interaction, because of the steric repulsion effect of PEG chains. Biodistribution studies with the 

above MePEGcyanoacrylate–hexadecylcyanoacrylate copolymer nanoparticles revealed that the 

nanoparticles exhibited long-circulating properties. 102 No effect of the degree of PEGylation on the 

plasma retention was observed because increasing the copolymer from 1:5 to 1:2 did not improve 

the nanoparticle circulation time. The 1:5 PEG–PHDCA copolymer probably already exhibited a 

PEG content that was able to determine a brush PEG configuration 103 and, thus, an optimal PEG 

density at the particle surface for the steric repulsion to occur efficiently. Together, the nanoparticles 

showed low liver accumulation and high spleen uptake, representing the possibility of 

targeting drugs to this tissue. More recently, a rapidly biodegradable copolymer poly(methoxyethylene 

glycol-cyanoacrylate-co-n-hexadecylcyanoacrylate) has been developed for the preparation of 

stealth nanoparticles. 102,104 Similarly, Ping et al. 105 synthesized poly(methoxyethyleneglycol 

cyanoacrylate-co-n-hexadecyl cyanoacrylate) (PEGylated PHDCA) nanoparticles loaded with an 

antitumor agent salvicine. The nanoparticles showed a significant initial burst of salvicine followed 

by a sustained release in 7.4 pH phosphate-buffered saline, thus suggesting its usefulness in 

tumor therapy. 

15.2.5 TOXICITY OF POLYALKYL CYANOACRYLATE NANOPARTICLES 


Toxicity of PACA nanoparticles varies depending on the alkyl chain length in the polymer. Incubation 

of hepatocytes with 1% poly(methyl cyanoacrylate) nanoparticles resulted in complete 

perforation of cell membranes. 106,107 However, PIBCA nanoparticles showed no signs of toxicity 

at the same concentration. This toxicity is probably due to the presence of nanoparticle degradation 

products in the cytoplasm following their phagocytosis. The low toxicity of isobutyl product is due 

to its slower degradation than the methyl product. The toxicity and safety data on PBCA nanoparticles 

is somewhat contradictory. Kante et al. 106 found LD50 of 230 mg/kg of PBCA 

nanoparticles injected intravenously in mice. Olivier et al. 108 observed a 30–40% mortality rate 

in mice after the injection of 166 mg/kg of PBCA nanoparticles; some animals died even at low 



doses. Meanwhile, Simeonova et al., 109 in an investigation of the immunomodulating properties of 

PBCA nanoparticles in mice, did not report any signs of toxicity at doses as high as 200 mg/kg. A 

much higher LD 50 of 585 mg/kg was found for nanoparticles made of a similar polymer PHCA. 110 

Gelperina et al. 111 reported no mortality in rats even after administering up to 400 mg/kg of empty 

PBCA nanoparticles. 

The implantation into the tissue of monomers, which polymerize within the body in orthopedic 

surgery and percutaneous minimally invasive radiological procedures, was performed earlier. The 

polymerization was mostly exothermic, and is associated with the toxicity to the tissue. 112,113 The 

use of n-butylcyanoacrylate in the endovascular treatment of arteriovenous malformations has been 

well documented. 113,114 Implantation of n-butyl-2-cyanoacrylate (NBCA) has resulted in a good 

occlusive effect on the parenchyma of highly vascular tumors, with extensive intralesional necrosis. 

The antitumor effect of the NBCA at the implant surface on adjacent tumors appeared limited. 115 

Histopathological examination of a surgically removed malignant pelvic para-ganglioma after a 

percutaneous injection of NBCA, three days after the NBCA injection, revealed variable degrees of 

tumoral necrosis. 116 Areas surrounding small volumes of NBCA at the periphery of the tumor 

showed practically no necrosis, whereas regions inside the tumor plastified with NBCA presented 

greater degrees of necrosis, especially toward the center of the tumoral mass (Figure 15.3). 

However, no necrosis was found at a macroscopic level outside the implant, indicating that the 

NBCA allowed for the destruction of tumor tissue within the heavily plastified areas. The carcinogenic 

potential of cyanoacrylate adhesive and isobutyl-2-cyanoacrylate were also reported 

respectively by Matsumoto 117 and Samson and Marshal. 118 Induction of malignant fibrous histiocytoma 

in female Fisher rats by implantation of foreign bodies was studied by Hatanaka et al. 119 

Five kinds of foreign bodies (silicone, cellulose, polyvinyl chloride, and zirconia and alkyl-alpha 

cyanoacrylate) were implanted into the subcutaneous tissue of female Fisher rats. The tumors 

developed in almost all the cases was composed of a mixture of cells that resembled fibroblast, 

Brain concentration (μg/g) 

7 

6 

5 

4 

3 

2 

1 

0 

–60 

0 

60 120 180 240 300 360 420 480 

Time [min] 

FIGURE 15.3 Uptake by rat brain of i.v. injected doxorubicin loaded to polysorbate 80-coated nanoparticles. 

Doxorubicin levels in rat brain (mg/g) versus time (min). C Doxorubicin (5 mg/kg) in saline. $ Doxorubicin 

(5 mg/kg) bound to nanoparticles overcoated with 1% polysorbate 80. (From Gulyaev, A. E. et al., Pharm. 

Res., 16, 1564, 1999. With permission.) 


266 

Golgi apparatus, histiocytes, myofibroblasts, and immature mesenchymal cells. A rat malignant 

fibrous histiocytoma transplanted into the subcutaneous tissue of the syngeneic female Fisher rats 

grew and metastasized to the lungs. Canter et al. 120 reported the consecutive pathological findings 

of a patient who underwent surgery for facial hemangioma after percutaneous injection of NBCA 

for devascularization of a lesion. The patient underwent additional surgery one month and six 

months after the initial operation for the removal of the residual NBCA cast from the injection 

site. Acute inflammatory findings after injection of NBCA and the development of a chronic 

granulomatous foreign-body reaction support the histological findings of experimental animal 

studies and postmortem examinations of humans. Though the alkylcyanoacrylates of smaller 

alkyl chain lengths exhibited considerable toxicity, the alkylcyanoacrylates of larger chain 

lengths such as isohexylcyanoacrylate nanoparticles resulted in negligible toxicity 121 and drove 

them to their clinical trials. 

15.2.5.1 Toxicity of Drugs Associated with Poly(Alkyl Cyanoacrylate) Nanoparticles 

The study of the acute toxicity of doxorubicin associated with polysorbate 80-coated nanoparticles 

in healthy rats and rats with intracranially implanted 101/8 glioblastoma was reported by Gelperina 

et al. 111 Single intravenous administration of empty PBCA nanoparticles in the dose range of 100– 

400 mg/kg did not cause mortality within the period of observation and also did not affect body 

weight or the weight of the internal organs. Association of doxorubicin with PBCA nanoparticles 

did not produce significant changes of quantitative parameters of acute toxicity of the antitumor 

agent. Efficacy and toxicity of mitoxantrone-loaded PACA nanoparticles were compared with a 

drug solution and with a mitoxantrone-liposome formulation. 89 Furthermore, the influence of an 

additional coating surfactant, poloxamine 1508, which has been shown to change the body distribution 

of other polymeric nanoparticles, was investigated. PACA nanoparticles led to a significant 

reduction in tumor volume of B-16 melanoma in mice compared to liposomes. However, neither 

nanoparticles nor liposomes had overcome the mitoxantrone-associated leucocytopenia. 

In a study by Gibaud et al. 122 involving the determination of the myelosuppressive effects of 

free and PACA-bound doxorubicin in mice in vivo, the alkyl cyanoacrylate nanoparticle-bound 

doxorubicin showed the highest myelosuppressive effect. The total and differential counts of blood, 

bone marrow, and spleen cells, as well as the number of granulocyte progenitors (CFU-GM), was 

determined by culture after the intravenous injection of 11 mg/kg body weight of doxorubicin, 

either free or bound to PIBCA or PIHCA nanoparticles. Doxorubcin bound to PIHCA nanoparticles 

showed the highest and longest myelosuppressive effects, which correlated well with a high 

concentration of the drug in the bone marrow and spleen. Furthermore, the PIHCA nanoparticles 

induced the release of colony-stimulating factors, which might account for the observed increase in 

the toxic effects of doxorubicin on bone marrow progenitors. 

15.2.6 DRUG DELIVERY IN CANCER WITH POLY(ALKYL CYANOACRYLATE) NANOPARTICLES 

15.2.6.1 Targeting to Cancer Cells and Tissues 


Increase in the antitumoral activity of many anti-cancer drugs was also observed when loaded in 

polyisohexylcyanoacrylate nanoparticles. A study 123 conducted with doxorubicin-loaded polyisohexylcyanoacrylate 

nanoparticles administered i.v. to the C57BL/6 mice consisting of metastases 

induced by i.v. inoculation of reticulosarcoma M5076 cell suspension showed a dramatic increase 

in the antitumoral activity of the drug. Doxorubicin measurements in healthy hepatic or neoplastic 

tissue, carried out with histological examinations using TEM, demonstrated that the hepatic tissue 

is an efficient reservoir of the drug when it was injected in association with nanoparticles. Accumulation 

of doxorubicin-associated nanoparticles suggests that the Kupffer cells created a gradient of 

drug concentration for a massive and prolonged diffusion of the free drug toward the neoplastic 

tissue. The negatively charged mitoxantrone PBCA nanoparticles (DHAQ-PBCA-NP) prepared by 



EP method 124 following intravenous administration were concentrated mainly in liver tumors. The 

accumulation of DHAQ-PBCA-NP was higher in the liver tumors than in the liver tissue. Yang et 

al. 125 compared antitumor activity of mitoxantrone (DHAQ) and mitoxantrone-polybutyl cyanoacrylate-nanosphere 

(DHAQ-PBCA-NS) against experimental liver tumor H22 in mice. The results 

of antitumor activity determination of both drugs with different treatment schedules showed 

DHAQ-PBCA-NS to present higher activity than DHAQ; DHAQ-PBCA-NS is possessed of 

liver-targeting property. 

Soma et al. 126 demonstrated that the Doxorubicin-loaded PACA nanoparticles are more 

efficient than free drugs in mice bearing hepatic metastasis of the M5076 tumor. High phagocytic 

activity of Kupffer cells in the liver played the role of drug reservoir after nanoparticle phagocytosis. 

The study also assessed the role of macrophages in mediating the cytotoxicity of doxorubicin-loaded 

nanoparticles on M5076 cells. After the phagocytosis of the doxorubicin-loaded nanoparticles, 

J774.A1 cells were able to release the active drug, allowing it to exert its cytotoxicity against 

M5076 cells. Intracellular accumulation and DNA binding of doxorubicin encapsulated in PIHCA 

nanospheres and of daunorubicin bound to polyglutamic acid (DGA) in comparison with free Dox 

and daunorubicin was studied by Bogush et al. 127 using methods involving interaction between an 

anthracycline and cellular DNA to understand how the drug reaches its nuclear targets. The results 

showed that the intracellular accumulation and DNA binding of Dox-loaded nanospheres and DGA 

are reduced by 30–40% in comparison with those obtained for free doxorubicin and 

daunorubicin, respectively. 

Studies by Bennis et al. 128 on the cytotoxicity and accumulation of doxorubicin encapsulated in 

PIHCA nanospheres in a model of doxorubicin-resistant rat glioblastoma variants revealed that the 

cytotoxicity differed with the degree of resistance to this drug. The nanoparticle associated Dox was 

more cytotoxic than the free Dox, whereas co-administration of drug-free nanoparticles with free 

Dox did not significantly modify the cytotoxicity of Dox. However, Simeonova and Antcheva 129 

reported that the cytotoxic activities of drug-free PBCA nanoparticles free of vinblastine, vinblastine-loaded 

nanoparticles (by incorporation and adsorption processes), and a mixture of vinblastinefree 

and drug-free PBCA nanoparticles, when compared in vitro on human erythroleukemic K-562 

cells, showed enhanced cytotoxicity when vinblastine was either adsorbed on PBCA or mixed with 

them rather than free. When vinblastine was incorporated into the polymer matrix of nanoparticles, 

a lag period and a postponed cytotoxic effect on K-562 cells was observed. Preclinical studies 121 in 

21 patients with refractory solid tumors also confirmed an increase in doxorubicin cytotoxicity and 

a decrease in cardiotoxicity when incorporated into biodegradable PIHCA by modifying tissue 

distribution. In clinical studies 130 conducted on 21 patients with Stages III–IV maxillary tumors, 

antitumor drugs deposited with a cyanoacrylate adhesive solution showed immediate and late 

results of therapy. This evidence of deposition of the drug and its regulated elimination from the 

composition helps create a high concentration of the agent at the operation site where the tumor 

elements remained after surgery. 

Pan et al. 131 demonstrated the higher efficacy of 5-FU when loaded into PBCA nanocapsules 

with a diameter of 338 nm, compared to the free drug in S180, HepS, and EAC tumors. 

The in vitro and in vivo antihepatoma effects of the liver-targeted drug delivery system lyophilized 

aclacinomycin-A loaded polyisobutylcyanoacrylate nanoparticle (ACM-IBC-NP) was found 

to be significantly concentration dependent. 132 The antihepatoma activity of lyophilized 

ACM-IBC-NP was higher than that of the aclacinomycin-A. Similarly, actinomycin-D absorbed 

into polymethylcyanoacrylate nanoparticles showed increased efficiency against an experimental 

tumor in comparison to the free drug. 133 

The antitumor efficacy of a medical-grade cyanoacrylate adhesive MK-8-based composition 

with prospidin has been studied by Chissov et al., 134 and the kinetics of the drug elimination have 

been followed up in a rat model. The findings showed a prolongation of the relapse-free period and 

a longer survival of the experimental animals. Clinical trials of the composition carried out in 70 

patients with esophageal carcinomas showed local disseminated forms in 52 patients. Application 


268 

of the adhesive composition has not increased the rate of postoperative complications; in some 

cases, the life span of patients after palliative surgery was considerably longer. In a study by Pérez 

et al. 135 using an N-butyl-2-cyano acrylate-based tissue adhesive, Tisuacryl, as a nonsuture method 

for closing wounds in oral surgery involving apicectomy, molar extractions, and mucogingival 

grafting, the adhesive was well tolerated by the tissue and permitted immediate hemostasis and 

normal healing of incisions. 

Association of 5-FU to PBCA nanoparticles also showed enhanced efficacy against Crocker 

sarcoma S 180 and a higher toxicity of the drug, as measured by induced leukopenia, body weight 

loss, and premature death. 31 The efficacy was further increased by an increase in the polymerto-drug 

ratio. The nanoparticles yielded a prolonged persistence of the 5-FU in all organs examined, 

including the tumor. 

15.2.6.2 Drug Delivery to Resistant Tumor Cells 


Several cancers show resistance to conventional chemotherapy. Resistance to cancer chemotherapy 

involves both altered drug activity at the designated target and modified intra-tumor pharmacokinetics 

(e.g., uptake and metabolism). The membrane transporter P-glycoprotein plays a major role 

in pharmacokinetic resistance by preventing sufficient intracellular accumulation of several anticancer 

agents. Inhibiting the P-glycoprotein has great potential to restore chemotherapeutic effectiveness. 

136 PACA nanoparticles were found to overcome multi-drug resistance (MDR) phenomena 

at both the cellular and noncellular level. 137,138 

Doxorubicin is a well known P-glycoprotein substrate. The resistant cells treated with doxorubicin-loaded 

PACA nanoparticles showed much higher sensitivity to the drug relative to the free 

doxorubicin. 139 Dox and Dox-loaded PIHCA nanoparticles were tested in two human tumor cell 

lines, K562 and MCF7, at increasing concentrations. The cell lines were more resistant to the free 

Dox than to the Dox-loaded nanoparticles. However, the MCF7 sublines showed a higher level of 

resistance to Dox-loaded nanoparticles than that of the free Dox. Different levels of overexpression 

of several genes involved in drug resistance (MDR1, MRP1, BCRP, and TOP2alpha) occurred in 

the resistant variants. MDR1 gene overexpression was consistently higher in free-Dox selected cells 

than in Dox-loaded nanoparticle-selected cells; this was the reverse for the BCRP gene. Overexpression 

of the MRP1 and TOP2alpha genes was also observed in the selected variants. The 

results indicate that several mechanisms may be involved in the acquisition of drug resistance and 

that drug encapsulation markedly alters or delays these processes. 140 

Further experiments, such as uptake studies in the presence of cytochalasin B or efflux studies, 

indicated a possible mechanism of nanoparticle–cell interaction. The degradation of the drug carrier 

was also shown to play a key role in the mechanism of action. The formation of an ion pair by the 

polycyanoacrylic acid resulting from the nanoparticle degradation with doxorubicin 141 has been 

considered as one of the possibilities (Figure 15.4). It was hypothesized that the nanoparticles 

adhere to the cell surface, followed by the release of doxorubicin and nanoparticle degradation 

products, which combine as an ion pair able to cross the cell membrane without being recognized 

by the P-glycoprotein (Figure 15.5). 142,143 The ion pair was considered to form between the 

positively charged doxorubicin and the negatively charged polycyanoacrylic acid, resulting in 

the charge masking of the doxorubicin and thereby making it more lipophilic and able to cross 

the cell membrane. The formation of an ion pair of 2-cyano-2-butyl hexanoic acid (a PACA-like 

molecule) with the doxorubicin has been assessed using ion pair reverse-phase high-performance 

liquid chromatography. As the PACA possess high molecular weight (w1000) and their nanoparticles 

possess large polydispersity, a PACA-like molecule, i.e., 2-cyano-2-butyl hexanoic acid 

(molecular weight of 197.28 g/mol) was used to investigate the formation of the ion pair. The 

ion pair formation was found between the charged amino group of doxorubicin and the carboxyl 

group of 2-cyano-2-butyl hexanoic acid, leading to the overall decrease in charge of the drug and 

the increase in drug lipophilicity (Figure 15.6). The experimental results were justified based on two 



FIGURE 15.4 Histological section of resected paraganglioma after percutaneous NBCA injection. (a) Section 

close to the tumor surface showing NBCA (void) and viable tumor around the implant (H&E, 

original magnification X25). (b) High-power view of (a). Note absence of necrosis around the implant 

(H&E, original magnification X100). (c) Section in a deeper zone of the tumor. Note extensive necrosis 

(H&E, original magnification X25). (d) High-power view of (c). Note an acellular necrotic area (H&E, original 

magnification X100). (From San Millan Ruiz, D. et al., Bone, 25(Suppl. 2), 85S, 1999. With permission.) 

assumptions: (1) the 2-cyano-2-butyl hexanoic acid and PACAs behave similarly in terms of acidic 

properties, and (2) the organic and aqueous buffer mixtures used in the chromatographic procedure 

could be correlated with the physiological medium. 

In another experiment, 142 C6 rat glioblastoma cells—the sensitive, resistant C6 0.001 (continuously 

grown in 0.001 mg doxorubicin/mL medium) and the C6 0.5 cells (continuously grown in 

0.5 mg doxorubicin/ml medium)—and the drug/formulation were separated by a polycarbonate 

membrane (pore size 0.2 mm) (Figure 15.7). The C6 sensitive cells showed the same cytotoxicity 

in the presence or absence of the membrane. While the doxorubicin-loaded PIHCA nanoparticles 

were found to overcome the resistance of C6 0.001 and C6 0.5 compared to the free drug and free 

drug C blank nanoparticles when in direct contact with the drug, a 4-fold and 5-fold resistance, 

respectively, was maintained when separated by the membrane. This clearly indicates the necessity 

of contact of the drug-loaded nanoparticles with the cell membranes to overcome the MDR in 

resistant cells. To confirm the possible interference of the nanoparticle degradation products with 

P-glycoprotein, the inhibition study of tritiated azidopine binding to the P-glycoprotein enriched 

membranes of C6 0.5 cells was performed. The blank nanoparticles could not inhibit the azidopine 

binding even after spontaneous degradation, whereas both the free doxorubicin and its nanoparticle 

formulation decreased the binding, indicating that the degradation products of nanoparticles do not 

have a direct role in P-glycoprotein (P-gp) inhibition and that it is the combination of both the 


270 

nucleus 

nanoparticle 

P-gp 

(a) (b) 

Nanoparticle in the 

course of degradation 

nucleus 

Ion pair 

Doxorubicin 

FIGURE 15.5 Hypothesis about the mechanism of action of PACA nanoparticles to overcome MDR at the 

cellular level. Drug-loaded nanoparticles are not endocytosed by the resistant cells (a) but adhere to the cell 

surface where they degrade and simultaneously release degradation products and the drug (b). The degradation 

products and the drug form ion pairs (c) that can penetrate the cells without being recognized by the P-gp and, 

by this means, increase the intracellular concentration of the anti-cancer drug in the resistant cells. (From 

Vauthier, C. et al., J. Control. Release, 93, 151, 2003. With permission.) 

doxorubicin and the PACA in the nanoparticle formulation that plays a role in 

P-glycoprotein inhibition. 

Among the various strategies implemented for bypassing the MDR in cancer cells (such as the 

use of ribozymes, 144 oligonucleotides, 145 and chemosensitizing compounds), the co-administration 

of chemosensitizing compounds with the anti-cancer agents has been widely investigated. 146 

Of such compounds, cyclosporine A, a potent immunosuppressive agent, was shown to prevent 

MDR in resistant cells in culture 147 and exhibited promising clinical results in refractory 

leukemia. 148 Cyclosporine A is believed to bind directly to P-gp, thereby inhibiting the efflux 

pump of the cytotoxic agents, resulting in their higher intracellular accumulation. Co-encapsulation 

of cyclosporine A and doxorubicin into polyisobutyl cyanoacrylate nanoparticles has resulted in 

a significant improvement in the reversion of MDR in P388-resistant cells compared to 

the mixed solution of doxorubicin and cyclosporine A and the nanoparticle-loaded doxorubicin 

incubated with free cyclosporine A. 149 The nanoparticles target the loaded drugs close to the cell 

membrane, providing the higher drug concentrations at the cell surface, 150 and this could be the 

reason for the higher efficacy of drugs loaded into the nanoparticles compared to those in 

solution form. 

Co-administration of cyclosporine A with anti-cancer agents was shown to overcome the multidrug 

resistance in resistant cell lines. 147 To understand the contribution of macrophages in delivering 

the doxorubicin-loaded nanoparticles to the tumor cells, a co-culture of sensitive or resistant 

P388 tumor cells and macrophage monocyte cells J774.A1 were incubated with the free doxorubicin, 

doxorubicin-loaded isobutylcyanoacrylate nanoparticles, and a combined cyclosporine A 

and doxorubicin-loaded nanoparticle formulation. 151 When incubated directly with tumor cells, 



+ 

(c)


(a) 

+ + 

− −− 

Doxorubicin 

O 

OCH3 O 

H 3 C 

+ 

OH 

OH H O 

OH 

O 

NH 3+ 

+ + + 

− −− 

+ + 

O 

C CH2OH OH 

O 

−−− 

CN 

− 

+ + 

−− 

HO 

CH 3 CH 3 ONa 

Butylbromide 

OCH2CH3 (1) 

O 

OH 

(1) Ethyl cyanoacetate 

(2) Ethyl (2-cyano 2-butyl) hexanoate 

(3) (2-cyano 2-butyl) hexanoic acid 

(b) 

CN 

+ 

−−− 

PACA degradation product 

−−− 

CH 2 

C 4 H 9 

C 4 H 9 

(3) 

O 

CN 

C 

C 

O − 

H 

O 

n=3 

CN 

C 4 H 9 

C 4 H 9 

OCH 2 CH 3 (2) 

FIGURE 15.6 2-Cyano-2-butyl hexanoic acid and doxorubicin. (a) Hypothesis on the biological action of 

PACA ion pair formation with doxorubicin. (b) Structure of the hypothetic ion pair formation and synthesis of 

the 2-cyano-2-butyl hexanoic acid (From Pepin, X. et al., J. Chromatogr. B, 702, 181, 1997. With permission.) 

medium only 

Cell monolayer 

KOH 

drug + medium Insert 

Well of a 24-well 

microplate 

polycarbonate membrane, 0.20 mm 

FIGURE 15.7 Schematic representation of the experimental device used for separating cell and drug compartments 

by a polycarbonate membrane (pore size 0.20 km). (From Hu, Y.-P. et al., Cancer Chemother. 

Pharmacol., 37, 556, 1996. With permission.) 


272 

both the doxorubicin and the doxorubicin-loaded nanoparticle formulation were equally cytotoxic. 

However, in co-culture, the doxorubicin-loaded nanoparticle formulation was more effective than 

the free doxorubicin. In co-culture, whereas the nanoparticle-loaded doxorubicin enhanced the 

cytotoxicity compared to the free drug, it has partly overcome the multidrug resistance in resistant 

cells. The addition of cyclosporine A improved the cytotoxicity of doxorubicin irrespective of the 

formulation, but it could not overcome the MDR. However, the efficacy of cyclosporine A in 

overcoming the MDR was higher in the case of the combined nanoparticle formulation with the 

doxorubicin, compared to the effect seen with the mixed nanoparticles of cyclosporine A and the 

doxorubicin prepared separately. Such a combinatorial nanoparticulate system would have both 

enhanced efficacy against the tumor and be able to overcome the adverse effects associated with the 

administration in the free form. 

Doxorubicin-14-O-hemiadipate (a 14-O-hemiadipate of doxorubicin) (H-Dox) was shown to 

partially circumvent the P-gp in MCF7/R human breast carcinoma and P388/R murine leukemia 

cell lines. 152 The comparative studies of the accumulation, retention, and intracellular localization 

of H-Dox and Dox in Dox-sensitive murine leukemia cell line P388/S and its resistant variant, Pgppositive 

drug-resistant P388/R subline, in the presence or absence of cyclosporine A, showed a 

38-fold decrease in Dox accumulation but only a 5-fold decrease in H-Dox accumulation, indicating 

a greater than 7-fold increase in H-Dox buildup in resistant cells. Coincubation with 

cyclosporine A resulted in a 54-fold increase in Dox accumulation and only a 5-fold increase in 

H-Dox uptake in P388/R cells, restoring the doxorubicin levels in P388/R to 100% of that found in 

P388/S cells. Once internalized by the resistant cells, H-Dox was retained better than Dox, regardless 

of the presence or absence of CsA. The H-Dox was localized in the nuclei of the P388/R cells 

but not the Dox. 153 Coadministration of chemosensitizers like the cyclosporine analogue SDZ PSC 

833 [(3 0 -keto-Bmt1)-(Val2)-cyclosporin] (PSC 833) and anti-cancer drugs has been shown to 

possess powerful chemosensitization properties in vitro, in addition to being intrinsically 

nontoxic. 154 

Complete reversion of drug resistance in vitro was demonstrated by Cuvier et al. 137 with 

Doxorubicin-loaded PACA nanoparticles in terms of cell growth inhibition comparable with that 

obtained with sensitive cells exposed to free Dox. In vivo experiments showed significantly 

prolonged survival of P388-Adr-resistant leukemia bearing mice that had previously received 

P388-Adr-R cells by i.p. injections of Dox-NS, whereas free Dox injection was ineffective 

against this rapidly growing tumor. Intracellular concentration and the cytoplasmic and nuclear 

distribution of Dox as measured by laser microspectrofluorometry (LMSF) showed a similar 

accumulation and distribution of Dox, regardless of the form of Dox delivery in the sensitive 

cells, whereas Dox-NS accumulated the same amount in resistant cells. 

Use of folate receptor-specific drug conjugates and their nanoparticles/liposomes forms is 

another strategy for bypassing multi-drug resistant efflux pumps. Folic acid, attached to polyethyleneglycol-derivatized, 

distearoyl-phosphatidylethanolamine, was used by Goren et al. 155 to target 

liposomes to folate receptor (FR)-overexpressing tumor cells in vitro. Additional experiments with 

DOX-loaded, folate-targeted liposomes (FTLs) indicate that liposomal DOX is rapidly internalized, 

released in the cytoplasmic compartment, and, shortly thereafter, detected in the nucleus, the entire 

process lasting 1–2 h. FR-mediated cell uptake of targeted liposomal DOX into a multi-drug 

resistant subline of M109-HiFR cells (M109R-HiFR) was unaffected by P-glycoprotein-mediated 

drug efflux, in sharp contrast to the uptake of free DOX. 

15.2.6.3 Drug Delivery to Hepatocellular Carcinomas 


Hepatocellular carcinomas (HCCs) are generally identified clinically at an advanced stage, 

usually in combination with cirrhosis. Though optimal, surgical resection is associated with 

a high rate of recurrence. Approaches to prevent recurrence, like chemoembolization after 

surgery, have not proven to be beneficial. Both doxorubicin and cisplatin are frequently 



used in the treatment of these carcinomas, but the overall response rates are low and neither 

approach seems to prolong the survival. Newly emerging agents with promising results include 

90 

Y microspheres, antiangiogenesis agents, inhibitors of growth factors and their receptors, and 

K vitamins. 156 

HCC is known to be chemoresistant to anti-cancer drugs due to the expression of multi-drug 

resistant (MDR) transporters. Doxorubicin loaded into the PACA nanoparticles enhanced the 

efficiency of treatment in the M5076 murine hepatic tumor metastasis in mice. 91 After intravenous 

injection, the doxorubicin-loaded PBCA nanoparticles resulted in higher hepatic concentrations in 

tumor-bearing mice. Histology of the mouse liver after the intravenous injection of the nanoparticle 

formulation revealed greater accumulation of nanoparticles in the Kupfer cells but not in the tumor 

cells. 123 This was supposed to be the result of macrophage uptake and subsequent capture by the 

Kupfer cells, of the nanoparticle formulation owing to the hydrophobicity of the nanoparticles, and 

then the drug release close to the tumor cells in the liver. To better understand the role of macrophages 

in delivering the nanoparticle-bound doxorubicin to the tumor cells, a co-culture system was 

developed consisting of two compartments separated by a porous polyester membrane (Transwell 

clear insert). 126 The results were compared with the control experiments involving direct incubation 

of drug/formulations with the tumor cells. The M5076 cells were seeded into the lower 

compartment, whereas the J774.A1 macrophage monocyte cells were placed on the top of the 

membrane. The doxorubicin and doxorubicin-loaded isobutylcyanoacrylate nanoparticle 

formulation were added to the upper compartment containing macrophages. The doxorubicin 

and nanoparticle formulation showed a 5-fold increase in IC50 of M5076 tumor cells in the 

co-culture system. In some groups of experiments, the recombinant mouse interferon-g (IFNg) 

was added to the macrophage compartment before the addition of the formulation to activate 

the macrophages. This would result in the release of cytotoxic factors such as IFN-a and nitric 

oxide. 157 Doxorubicin and nanoparticle formulation were found to be more cytotoxic when the 

macrophages were activated by IFN-g. In fact, this would probably be a result of a synergistic 

effect of the cytotoxic factors, especially nitric oxide released by the activated macrophages and 

the nanoparticle formulation because the role of antitumor factors in the inhibition of tumors both 

in vitro and in vivo has been well documented. 158,159 This co-culture system gives a clear 

indication of the role that macrophages play after intravenous injection in enhancing the 

cytotoxicity of the doxorubicin-loaded PACA nanoparticle formulation and its efficacy in 

liver tumor. 

Recently, Barraud et al. 160 compared the antitumor efficacy of Dox-loaded PIHCA nanoparticles 

with free doxorubicin, both in vitro and in vivo, in HCC-bearing transgenic mice 

overexpressing the mdr1 and mdr3 genes. The 50% inhibition concentration (IC50) of doxorubicin 

in vitro on different human hepatoma cell lines decreased with the nanoparticle formulation, 

compared to that of free Dox. Dox-loaded nanoparticles showed an enhanced MDR reversal in 

these cells. 

Mitoxantrone-loaded PBCA nanoparticles showed higher tumor concentration after intravenous 

injection in the B-16 melanoma-bearing mice. 89 These nanoparticles have also shown higher 

efficacy in the treatment of B-16 melanoma. 28 The efficacy of mitoxantrone-loaded PBCA nanoparticles 

after intravenous administration was also tested in heterotopic and orthotopically 

transplanted human HCC-bearing mice. 161 The efficacy of the mitoxantrone-loaded nanoparticles 

was compared with that of free mitoxantrone and doxorubicin. In orthotopically transplanted 

mice, all three preparations produced a similar effect. Microscopic examination revealed that the 

mitoxantrone-loaded nanoparticles showed enhanced antitumor activity in heterotopically transplanted 

mice, which was supported by a low tumor cell proliferation. The nanoparticles also did not 

show any signs of toxicity of the heart, spleen, lung, and kidney, indicating their effectiveness 

in tumor therapy and that the fact that they had overcome the adverse effects associated with the 

anti-cancer agents. 


274 

15.2.6.4 Drug Delivery to Brain Tumors 


Brain tumors are one of the most lethal forms of cancer. Many primary brain tumors are most 

resistant to chemotherapy, probably due to the presence of a tight BBB. 162 They are extremely 

difficult to treat due to a lack of therapeutic strategies capable of overcoming barriers for effective 

delivery of drugs to the brain. In fact, the vasculature of gliomas possess some special features: 

open endothelial gaps (inter-endothelial junctions and transendothelial channels having diameter of 

about 0.3 mm), fenestrations (5.5 nm maximum width), cytoplasmic vesicles such as caveolae (50– 

70 nm diameter), and vesicular vacuolar organelles (108G32 nm diameter). All of these features 

are due to the secretion of a vascular endothelial growth factor (VEGF) that causes the loss of the 

barrier function of BBB, which is essential for delivering drugs to the brain. 163 Nanoparticles are 

considered suitable for delivering drugs to the brain, but they should possess the following ideal 

characteristics: 164 (1) a nontoxic, biodegradable, and biocompatible nature; (2) a particle diameter 

less than 100 nm; (3) exhibit physical stability in the blood (no aggregation); (4) avoidance of the 

MPS (no opsonization) prolonged blood circulation time; (5) a BBB-targeted brain delivery 

(receptor-mediated transcytosis across brain capillary endothelial cells); (6) a scalable and costeffective 

manufacturing process; (7) be amenable to small molecules, peptides, proteins, or nucleic 

acids; (8) exhibit minimal nanoparticle excipient-induced drug alteration (chemical degradation/ 

alteration, protein denaturation); and (9) show possible modulation of drug release profiles. 

Extensive efforts have been made to develop novel strategies to overcome the obstacles for 

brain tumor drug delivery. 165 PACA nanoparticles enabled the delivery of a number of drugs, 

including doxorubicin, loperamide, tubocurarine, the NMDA receptor antagonist MRZ 2/576, 

and the peptides dalargin and kytorphin across the BBB after coating with surfactants. 166 

However, only the surfactants polysorbate (Tween) 20, 40, 60, and 80, as well as some poloxamers 

(Pluronic F68), can induce this uptake. The mechanism for the delivery across the BBB most likely 

is endocytosis via the LDL receptor by the endothelial cells lining the brain blood capillaries after 

the injection of nanoparticles into the blood stream. This endocytotic uptake seems to be mediated 

by the adsorption of apolipoprotein B and/or E from the blood. After this process, the nanoparticleassociated 

drug may be released into the cells that diffuse into the internal parts of brain or the 

particles may be transcytosed. 

Unfortunately, the conventional approaches resulted in only sub-therapeutic concentrations in 

the brain, due to the short plasma half-life of the compounds as well as the carrier system. Surface 

coating the carrier system with polyethylene glycol led to the enhancement of plasma residence 

time, due to the formation of a steric barrier preventing the interaction of the carrier with the plasma 

proteins (opsonins) and thus preventing the opsonization and further uptake by the MPS. 92,101 

PEG-coated-polyalkylcyanoacrylate nanoparticles consisting of the amphiphilic copolymer 

poly(PEGCA-co-hexadecyl cyanoacrylate) showed longer blood circulation time and better brain 

penetration than the nanoparticles coated with polysorbate 80 after intravenous injection in mice 

and rats. Unlike the polysorbate 80-coated nanoparticles, the penetration of poly(PEGCA-cohexadecyl 

cyanoacrylate) nanoparticles into the brain occurred without any modification of the 

BBB. The brain permeability of the nanoparticles further increased in pathological conditions, 

where the permeability of the BBB itself is modified. This has been demonstrated with rats 

bearing an experimental allergic encephalomyelitis after intravenous injection. 167 

PEG-coated hexadecylcyanoacrylate nanospheres synthesized by Brigger et al. 168 showed a 

significant accumulation after intravenous injection in gliosarcoma in rats. The radioactive [ 14 C]poly(hexadecylcyanoacrylate) 

([ 14 C]-PHDCA) and [ 14 C]-poly(MePEG2000cyanoacrylate-cohexadecylcyanoacrylate) 

1:4 (PEG–PHDCA) ([ 14 C]-PEG–PHDCA) nanospheres were prepared 

using the nanoprecipitation technique. This procedure involves the dissolution of the radioactive 

polymer in a warm acetone, followed by the addition of the acetonic solution into the aqueous 

phase, resulting in the precipitation of nanospheres. The [ 14 C]-PHDCA and [ 14 C]-PEG–PHDCA 

nanospheres possessed a mean diameter of about 150 nm. Intravenous injection of the 



[ 14 C]-PEG–PHDCA nanospheres into the rats inoculated intracerebrally with 9L gliosarcoma cells 

resulted in a long circulation time in the blood. Both [ 14 C]-PHDCA and [ 14 C]-PEG–PHDCA 

nanospheres were able to selectively extravasate the BBB and accumulated more in the gliosarcoma 

than in the peritumoral brain. However, this effect was significantly higher (3.1-times greater 

accumulation (p!0.05)) for [ 14 C]-PEG–PHDCA nanospheres than for the [ 14 C]-PHDCA nanospheres. 

Intravenous injection of a hydrophilic tracer [ 3 H]-sucrose to the tumor-bearing rats (seven 

days after the inoculation of tumor cells) to assess the permeability of the BBB showed that the 9L 

gliosarcoma and, to a lesser extent, the peritumoral brain region were hyperpermeable to sucrose, 

indicating a selective disruption of the BBB at the pathological site and slight edematous 

surrounding the brain. The permeability of the BBB was not observed at the intracerebral injection 

site of the control rats, indicating that the BBB alteration was only due to the presence of 

cancerous cells. 

15.2.6.5 Drug Delivery to Breast Cancer 

Resistance to cancer chemotherapy involves both altered drug activity at the designated target and 

modified intra-tumor pharmacokinetics (e.g., uptake and metabolism). Two proteins in particular— 

P-glycoprotein (P-gp) and MDR-associated protein (MRP2)—are responsible for MDR associated 

with a variety of cancers. 169 BCRP (breast cancer resistant protein) is another type of protein that 

appears to play a major role in the MDR phenotype of a specific human breast cancer. 170 Cancer 

cells overexpressing these proteins are resistant to several chemotherapeutic agents, including the 

anthracyclines, and show a reduced nuclear accumulation of anthracyclines. 171 The ability of 

PACA nanoparticles to overcome the MDR of cancer cells has been well documented. 137,172 

MRP1 is considered to be a transporter of some conjugated organic anions and drug-glutathione 

conjugates. 173 When delivered through PACA nanoparticles, doxorubicin forms an ion pair with 

the polycyanoacrylic acid and penetrates the cell membrane in resistant cells. 141 The efficacy of 

doxorubicin-loaded PIHCA nanoparticles was evaluated in resistant human breast adenocarcinoma 

cells, i.e., MCF7/VP and MCF7/doxorubicin overexpressing MRP1 and P-glycoprotein, respectively. 

171 The MCF7/VP and MCF7/doxorubicin were 34 and 47 times resistant to doxorubicin 

alone, and 2.5 and 7.7 times resistant to nanoparticle formulation, respectively, compared to MCF7 

cells. The nuclear accumulation of doxorubicin at 6 h after treatment was higher in MCF7 cells 

(133G22 and 120G41 mM when administered as free doxorubicin and nanoparticle formulation, 

respectively), followed by MCF7/VP (97G19 and 107G15 mM administered as free doxorubicin 

and nanoparticle formulation, respectively), and MCF7/doxorubicin (59G12 and 66G12 mM 

administered as free doxorubicin and nanoparticle formulation, respectively). At 12 h post incubation, 

the nuclear accumulation of doxorubicin in MCF7 cells did not increase significantly, 

whereas the accumulation was significantly higher in MCF7/VP cells (76G10 and 118G30 mM 

for free doxorubicin and nanoparticle formulation, respectively). At 18 h post incubation, the 

nuclear accumulation still increased in MCF7/VP cells (102G20 and 158G16 mM for free doxorubicin 

and nanoparticle formulation, respectively). However, the nuclear accumulation for MCF7/ 

doxorubicin cells did not vary much with time. The confocal microscopic studies indicated the 

localization of doxorubicin in the nucleus of MCF7 cells, whereas the localization was observed in 

the cytoplasm alone in MCF7/doxorubicin cells after treatment with free doxorubicin and nanoparticle 

formulation. In contrast, for MCF7/VP cells the localization was in both the nucleus and the 

cytoplasm in cells treated with nanoparticle formulation; it was only in the cytoplasm after 

doxorubicin treatment. 

It was hypothesized that after hydrolysis of the lateral ester bonds, the ion pair passively diffuse 

through the lipid bilayer of the cell membrane and some amount of the polycyanoacrylic acid 

(3–4 carboxylic group) could interact with glutathione and create a polycyanoacrylic acid–glutathione 

conjugate that would inhibit the efflux of MRP1–doxorubicin. The ion pair formation 

between doxorubicin and polyisohexylxyanoacrylate nanoparticles could also prevent the 


276 

formation of the doxorubicin–glutathione–S-conjugate and hence the MRP1-mediated transport of 

the drug. 171 Thus, the doxorubicin-loaded PACA nanoparticles overcome the MDR in cells overexpressing 

P-glycoprotein and MRP1 accompanied by the nuclear accumulation of doxorubicin. 

In support of the above results, Kisara et al. 174 observed that the two compounds buthionine 

sulfoximine (BSO) and cepharanthine (CE), which decrease the glutathione content in the MDR 

cells, showed greater reversal of multi-drug resistance to Dox in MDR cells. Treatment of the MDR 

cells with BSO resulted in an enhancement of the cytotoxic effect of Dox by 1.8-fold, whereas CE 

caused a greater reversal of drug resistance. BSO treatment also resulted in the decrease in GSH 

content of MDR cells compared to that of the sensitive ones. The combination of BSO with CE 

caused further potentiation of the antiproliferative effect of Dox in MDR cells. 

15.2.6.6 Drug Delivery to Lymphatic Carcinomas 

The lymphatic system is the physiological system that maintains the body water balance. It also 

controls certain immunological responses. At times, this system acts as a medium for the mestastais 

of cancer cells. Due to the peculiarity of the anatomy of the lymphatic interstitium, the achievement 

of drug localization into the lymphatic system is limited. Much effort has been made to achieve 

lymphatic targeting of drugs using colloidal carriers such as biodegradable nanoparticulate systems, 

including nanospheres, emulsions, and liposomes. 175 The major purpose of lymphatic targeting is 

(1) to provide an effective anti-cancer chemotherapy that will prevent the metastasis of tumor cells 

by accumulating the drug in the regional lymph node, and (2) to enhance the localization of 

diagnostic agents in the regional lymph node to visualize the lymphatic vessels before surgery. 

In the early days of lymphatic-targeting research, some success was achieved in delivering the 

emulsions and liposomes to the lymphatic system after intramuscular 176 and subcutaneous routes of 

administration. 177 For effective penetration into the lymphatic interstitium, the carrier systems need 

to possess a smaller size, hydrophobicity, and a negative surface charge. 178 PIBCA nanocapsules 

showed enhanced lymphatic targeting and regional lymph node retention of 12-(9-anthroxy) stearic 

acid, a lipophilic model compound. 175 The route of administration of the delivery system has a 

greater contribution in the delivery of nanoparticles to the lymphatic tumor in mice. 179 In vivo 

studies conducted by Reddy et al. 74 on mice bearing Dalton’s Lymphoma tumor have showed that 

99m Tc-Dox-loaded PBCA nanoparticles exhibited enhanced accumulation in the lymph node, 

compared with the 99m Tc-Dox, after subcutaneous injection. A significantly high tumor uptake 

of DPBC nanoparticles (w13-fold higher at 48 h post injection) (P!0.001) was found, compared 

to free doxorubicin. The accumulation of the nanoparticles in the tumor increased with time, 

indicating their slow penetration from the injection site into the tumor. Such a nanoparticulate 

system would be advantageous in the lymphoma treatment, offering drug targeting and the 

prolonged availability of the drug to the tumor. 

15.2.6.7 Drug Delivery to Gastric Carcinomas 


Gastric carcinoma falls into the category of the difficult-to-treat tumors. Though the PACA nanoparticles 

are considered most promising in the tumor-treatment applications, their use in the 

delivery of anti-cancer agents to gastric tumors remains limited. 

Magnetically targeted systems can be classified in the new generation of drug carriers. 

The advantage of these carriers is their ability to minimize the uptake by the reticuloendothelial 

system. 180 Intra-arterial administration of these systems in the form of magnetic albumin microspheres 

181 and magnetic liposomes 182 has been shown to treat the tumors successfully under the 

influence of a strong magnetic field. Recently, Gao et al. 183 prepared aclacinomycin-loaded 

magnetic iron oxide–PBCA nanoparticles using the interfacial polymerization technique. The 

technique involved dispersion of the drug in diluted hydrochloric acid into magnetic fluid, and 

the successive addition of the mixture to a hexane containing span 80 and polysorbate 80. 



The buytlcyanoacrylate monomer was added into the above dispersion under stirring, and after 6 h 

the newly formed magnetic nanoparticles were separated using a magnet, washed with methanol 

and water, lyophilized, and sterilized by 60 Co irradiation (15 kGy). These magnetic particles were 

210 nm in diameter, and they revealed a core-shell structure in which the iron oxide core was 

covered by the polymer coat. The aclacinomycin-loaded nanoparticles were evaluated in mice 

implanted with gastric carcinoma near the right fore feet. Administration of drug-loaded magnetic 

nanoparticles by intravenous injection into the tumor-bearing mice implanted with a magnet in the 

tumor resulted in a higher tumor-inhibition rate than the free aclacinomycin. The nanoparticle 

formulation greatly reduced the tumor mass compared to the free aclacinomycin and the drugfree 

carrier. The drug-loaded magnetic nanoparticles showed similar activity to that of the drug-free 

magnetic nanoparticles in vitro in the absence of a magnetic field in the gastric cancer cell line 

MKN-45. They did, however, show enhanced antitumor activity in vivo in tumor-bearing mice. 

15.2.6.8 Chemoembolization Using Alkylcyanoacrylates 

HCC is the most common liver tumor, with heterogeneity in the tumor behavior and the underlying 

liver disease. Recent combinations such as cisplatin, interferon, Adriamycin, and 5-FU are extremely 

toxic and yield response rates of only 20%, with no survival advantage compared to supportive care 

alone. 184 Higher concentrations of cancer chemotherapeutic agents can be delivered directly to the 

HCC via the hepatic arterial route. Considering that this route is the major vascular supply of these 

tumors, an even larger number of papers have reported the experience of hepatic artery 

chemotherapy or hepatic artery chemoembolization (TACE) with single agents, or with a dizzying 

combination of agents, and at doses not replicated by any two institutions. Loewe et al. 185 evaluated 

the potential of transarterial permanent embolization with the use of a mixture of cyanoacrylate and 

lipiodol for the treatment of unresectable primary HCC. Loewe et al. 186 used NBCA for hepatic 

artery embolization for the treatment of small-bowel neuroendocrine metastases to the liver. The 

results revealed that the permanent embolization of hepatic arteries as part of a multimodality 

treatment protocol is beneficial in long-term follow-up for patients with metastasized small-bowel 

neuroendocrine tumors. The use of cyanoacrylate is safe and effective as an embolic agent. 

Transarterial embolization (TAE) with the use of microspheres and Lipiodol and cyanoacrylate 

for unresectable HCC is a feasible treatment modality. A retrospective analysis of 46 patients with 

histologically confirmed HCC was made who were treated with TAE of the hepatic arteries. 187 To 

induce permanent embolization, the microspheres (Embosphere; 100–700 mm) and a mixture of 

ethiodized oil (Lipiodol Ultrafluide) with cyanoacrylate (Glubran) was administrated. No patient 

died during embolization or within the first 24 h. Severe procedure-related complications were 

observed in 2 patients. At the time of the analysis, 38 of 46 patients were alive. The 180-, 360-, 520- 

, and 700-day cumulative survival rates for the total study population were 80.6%, 70.7%, 70.7%, 

and 47.1%, respectively, with a median survival of 666 days. 

A procedure for effective and promising preoperative embolization of carotid body tumors was 

reported by Harman et al. 188 Ultrasound-guided direct percutaneous injection of n-butyl cyanoacrylate 

was given, and angiographic road map assistance was used for protection of parent arteries 

during the injection. After embolization, complete devascularization of the tumor was achieved 

without complications. The tumor was removed surgically with minimal blood loss. Transcatheter 

arterial embolization (TAE) of splanchnic arterial branches to allow continuous application of 

repeat hepatic arterial infusion chemotherapy (HAIC) was assessed. 189 One hundred and twentyeight 

patients with unresectable advanced liver cancer were implanted with a percutaneous port 

catheter system and TAE of splanchnic arteries with coils and/or NBCA. The recanalization rate 

between coil-embolized and NBCA- or NBCA-coil-embolized arteries, and frequency of heterogeneously 

poor distribution was compared between patients with single arteries and those with 

multiple hepatic arteries. The arteries once embolized with coils alone spontaneously recanalized at 

a significantly higher rate than those with NBCA. A hepatic artery embolization study carried out 


278 

by Loewe et al. 186 using NBCA and ethiodized oil for the treatment of small-bowel neuroendocrine 

metastases to the liver for the treatment of liver metastases from neuroendocrine 

small-bowel tumors also concluded that the use of cyanoacrylate as an embolic agent is safe 

and effective. 

The potential of transarterial permanent embolization with the use of a mixture of cyanoacrylate 

and lipiodol for treatment of unresectable primary HCC was also assessed by Loewe et 

al. 185 The study included 36 patients with histologically proven HCC who were treated with 

transarterial embolization of the hepatic arteries. The study indicated that TAE with use of 

cyanoacrylate and lipiodol for unresectable HCC is a feasible treatment modality. A similar 

study conducted by Berghammer et al. 190 confirmed the safety of the procedure with minimum 

side effects. Thus, it constitutes a valuable therapeutic option for patients with Okuda stage I 

and II HCC. 

A right gastric artery (RGA) embolization study to prevent acute gastric mucosal lesions caused 

by an influx of anti-cancer agents into the RGA in patients undergoing repeat HAIC was conducted 

on 217 patients with malignant hepatic tumors 191 using metallic coils and/or a mixture of n-butyl 

cyanoacrylate (n-BCA). RGA embolization was technically successful in the majority of patients 

(93%), with the lowest incidence of major complications. The clinical experience of Nadalini 

et al. 192 using isobutyl-2-cyanoacrylate vesical for embolization of the hypogastric arteries in 

cases of serious hemorrhage of the bladder and prostate is also reported. The effect was immediate 

results in the majority of cases, a decidedly positive outcome, especially considering the serious 

conditions of certain neoplastic patients. Isobutyl-2-cyanoacrylate suspension in Lipiodol was also 

used to treat percutaneous transcatheter embolization of the renal artery in clear cell carcinoma 193 

in nine patients. In most of the patients, the procedure was found to be palliative, with no complications 

attributed to the glue or to oil emboli. Preoperative embolization with isobutyl-2cyanoacrylate 

by means of an intra-arterial catheter during selective angiography was adopted 

by Carmignani et al. 194 in two cases of carcinoma of the kidney. 

Traditional preoperative embolization via a transarterial approach has proved beneficial, but it 

is often limited by complex vascular anatomy and unfavorable locations. 195 Paragangliomas, or 

glomus tumors, are the neoplasms of the head and neck. They are remarkably vascular, so surgical 

resection can be complicated by rapid and dramatic blood loss. 196 These tumors can develop in the 

middle ear (glomus tympanicum), the jugular foramen of the skull base (glomus jugulare), or the 

head and neck area (glomus caroticum, glomus vagale). Surgical removal of these tumors is also 

often associated with a significant intraoperative bleeding rate because of their vascular nature. 197–199 

Direct percutaneous injection of n-butyl cyanoacrylate resulted in the effective devascularization 

of craniofacial tumors 200 and the embolization of oral tumors. 201 However, the technique 

involved additional risks and was not widely adopted. In a study in human patients by Abud 

et al., 196 the presurgical devascularization was achieved by placing the diagnostic catheter in 

the common carotid artery to guide the puncture and perform the control angiography during 

and after the injection of the cyanoacrylate. The percutaneous devascularization of head 

and neck paragangliomas through the intralesional injection of cyanoacrylate resulted in effective 

devascularization. It could be a safe and effective technique for managing such 

clinical lesions. 

15.2.6.9 Delivery of Peptide Anti-Cancer Drugs 


Ras mutations exist in almost 20–30% of human tumors, and the ras oncogenes possess single-point 

mutations in the sequence coding for the active site of RAS protein at codons 12, 13b, and 61. 202 

These point mutations are good targets for the antisense oligonucleotides and can suppress the 

translation and targeted mutant mRNA. The mutated ras genes are involved in the cell proliferation 

and tumorigenicity. Antisense oligonucleotides (ODN) are molecules that are able to inhibit oncogene 

expression, being therefore potentially active for the treatment of viral infections or cancer. 



However, because of their poor stability in biological media and their weak intracellular 

penetration, colloidal drug carriers such as nanoparticles were employed for the delivery of 

oligonucleotides. 203 Association with nanoparticles protects the ODN against degradation and 

enables them to penetrate more easily into different types of cells. Thus, they were shown to 

improve the efficiency in inhibiting the proliferation of cells expressing the point-mutated Ha-ras 

gene. In vivo, the PACA nanoparticles were able to efficiently distribute the ODNs to the liver. As 

the ODNs have no affinity toward the PACA structure, the ion pair technique was adopted, using 

cationic surfactants such as cetyltrimethyl ammonium bromide or diethylaminoethyl dextran 

adsorbed onto the particle surface. 204,205 Spontaneous interpolymer complexation between 

cationic polyelectrolytes and DNA is known and is largely the result of cooperative electrostatic 

forces. 206 

ODNs adsorbed onto the PIHCA nanoparticles inhibited the Ha-Ras dependant tumor growth in 

mice. The ODN-nanoparticle formulation was tested in two cell lines, HBL100rasl and 

HBL100neo. 202 HBL100rasl is a clone obtained from the human mammary cell line. It expresses 

normal Ha-ras and Ha-ras carrying the G/U point mutation in codon 12 coding for Val instead of 

Gly at position 12. HBL100neo is a clone transformed only with the pSV2 vector, and it expresses 

only normal Ha-ras. Three 12-mer oligonucleotides—(1) an antisense oligonucleotide (AS-Val, 5 0 - 

CACCGACGGCGC-3 0 ) directed against and centered at the point mutation in codon 12 of the Ha-ras 

mRNA, (2) an antisense oligonucleotide (AS-Gly, 5 0 -CACCGCCGGCGC-3 0 )targetedtothe 

equivalent sequence of the normal Ha-ras mRNA, and (3) the 5 0 /3 0 inverted sequence of AS-Val 

that contains the same bases as the antisense sequence but in reverse orientation (INV-Val, 5 0 - 

CGCCGGAGCCAC-3 0 ). The inhibitory effect of AS-Val adsorbed on the nanoparticles was 

found to be at a concentration 100-times lower than that for the free AS-Val when tested on 

HBL100rasl. Cellular uptake studies using 32 P-labeling of free ODN and ODN adsorbed nanoparticle 

formulation revealed that the ODNs adsorbed onto the nanoparticles, incorporated at a lower 

rate than the free ODN but remained intact in the cells even after 72 h post incubation; the free ODN, 

in contrast, degraded after 3 h. Upon administration into the nude mice treated with HBL100rasl 

cells, the AS-Val nanoparticle formulation showed more potential tumor growth inhibition 

compared to the INV-Val nanoparticle formulation, indicating the sequence-dependant effect of 

ODN on the tumor-inhibition properties. 

The cationic surfactants used for the preparation of ODN–PACA nanoparticles may cause 

toxicity. In addition, the esterase-induced erosion of PACAs causes rapid desorption of ODNs in 

serum. 207 To overcome these drawbacks, the PACA nanocapsules with aqueous core have been 

designed. 208,209 These nanocapsules enhanced the protection and serum stability of the ODNs. 209 

Intratumoral injection of PACA nanoparticles containing Phosphorothioate oligonucleotides 

directed against EWS Fli-1 chimeric RNA led to a significant tumor growth inhibition compared 

to the free ODN in mice-bearing Ewing sarcoma. 

Peptides like Leu-enkephalin dalargin and the Met-enkephalin kyotorphin normally do not 

cross the BBB when given systemically. Transport of these neuropeptides across the BBB was 

achieved by adsorption onto the surface of PBCA nanoparticles and successive coating of the 

nanoparticles with polysorbate (Tween) 80. 210 

Birrenbach and Speiser 5 described the polymerization process for the preparation of hydrophilic 

micelles containing solubilized drug molecules including labile proteins in a colloidal 

aqueous system of dissolved monomers. The hydrocarbon medium constituted the outer phase. 

After secondary solubilization with the aid of selected surfactants, the polymerization of micelles 

under different conditions was performed. The entrapped tagged material (human 125I-immunoglobulin 

G) showed a stable fixation in the nanoparticles during long-term in vitro liberation trials. 

Nanoparticles were spherical in shape, 80 nm in diameter, and embedded with antigenic material 

(tetanus toxoid and human immunoglobulin G) for parenteral use. These preparations showed intact 

biological activity and high antibody production in animals. 


280 


Drug delivery to tumors has recently been the area of significant interest in the field of drug 

delivery. Nanoparticles, due to their many inherent advantages, gained an important place in this 

area of research. Though the alkylcyanoacrylates were initially used as tissue adhesives, the 

discovery of the simple anionic polymerization of alkylcyanoacrylates has renewed interest in 

their applicability to drug-delivery research. Poly(alkyl cyanoacrylates) with a higher alkyl chain 

length are relatively less toxic than those with a lower alkyl chain length. Several interesting studies 

were reported on the usefulness of poly(alkyl cyanoacrylates) for delivering anti-cancer drugs to 

different types of cancers, including breast cancer and gastric cancer. Interesting results were 

observed in the case of delivery of doxorubicin-loaded poly(alkyl cyanoacrylate) nanoparticles 

for the treatment of liver tumors. Though the initial results on the delivery of doxorubicin-loaded 

poly-butylcyanoacrylate nanoparticles to lymphoma are encouraging, more experimental evidence 

is needed to evaluate their efficacy on tumors. The PEG modification of alkylcyanoacrylates has 

improved the delivery of their nanoparticles to the brain. Cellular resistance to multiple drugs 

represents a major problem in cancer chemotherapy. This drug resistance may appear clinically 

either as a lack of tumor size reduction or as the occurrence of clinical relapse after an initial 

positive response to treatment with antitumor agents. The resistance mechanism can have different 

origins, either directly linked to specific mechanisms developed by the tumor tissue or connected to 

the more general problem of distribution of a drug toward its targeted tissue. The role of P-glycoprotein 

in decreasing the uptake of anti-cancer drugs in multi-drug resistant (MDR) tumors through 

efflux is well studied. Reports on the importance of nanoparticles in reversion of multi-drug 

resistance mediated by P-glycoprotein in MDR tumors reveal the potential of using PBCA nanoparticles 

in MDR tumors. A commercial application of such a strategy is on the way: doxorubicinloaded 

poly(alkyl cyanoacrylate) nanoparticles have reached phase-II clinical trials for the treatment 

of resistant tumors. Several reports have elucidated the mechanism of nanoparticle uptake by 

the cells, but the intracellular trafficking and fate of these nanoparticles in the cell remain unclear. 

Of late, poly(alkyl cyanoacrylate) nanoparticles have also been considered as a potential system for 

delivering proteins and peptide drugs to tumors. 

REFERENCES 


1. Fenton, R. G. and Longo, D. I., Cell biology of cancer, In Harrison’s Internal Medicine, Fauci, A. S., 

Braunwald, E., Isselbacher, K. J., Wilson, J. D., Martin, J. B., Kasper, D. L., Hauser, S. L., and 

Longo, D. I., Eds., vol. 1, 14 th ed., McGraw Hill, New York, p. 505, 1998. 

2. Reddy, L. H., Drug delivery to tumors: Recent strategies, J. Pharm. Pharmacol., 57, 1231, 2005. 

3. Williams, J. et al., Nanoparticle drug delivery system for intravenous delivery of topoisomerase 

inhibitors, J. Control. Release, 91, 167, 2003. 

4. De Jaeghere, F., Doelker, E., and Gurny, R., Nanoparticles, In Encyclopedia of Controlled Drug 

Delivery, Mathiowitz, E., Ed., Wiley, New York, p. 641, 1999. 

5. Birrenbach, B. and Speiser, P. P., Polymerized micelles and their use as adjuvants in immunology, 

J. Pharm. Sci., 65, 1763, 1976. 

6. Rhodes, C. T., Wai, K., and Banker, G. S., Molecular scale drug entrapment as a precise method of 

controlled drug release. II. Facilitated drug entrapment to polymeric colloidal dispersions, J. Pharm. 

Sci., 59, 1578, 1970. 

7. Boylan, J. C. and Banker, G. S., Molecular scale drug entrapment as a precise method of controlled 

drug release. IV. Entrapment of anionic drugs by polymer gelation, J. Pharm. Sci., 62, 1177, 1973. 

8. Langer, R., Biomaterials in drug delivery and tissue engineering: one laboratory’s experience, Acc. 

Chem. Res., 33, 94, 2000. 

9. Lanza, R. P., Langer, R., and Chick, W. L., Principles of Tissue Engineering, Academic Press, 

Austin, TX, p. 405, 1997. 



10. Gilbert, R. G., Emulsion Polymerization: A Mechanistic Approach, Academic Press, London, p. 30, 

1995. 

11. Bovey, F. A. et al., Emulsion Polymerization, vol. 1, Interscience Publishers, New York, p. 15, 1955. 

12. Ballard, M. J., Napper, D. H., and Gilbert, R. G., Kinetics of emulsion polymerization of methyl 

methacrylate. Polymer chemistry special issue, J. Polym. Sci., 22, 3225, 1984. 

13. Ljungstedt, I., Ekman, B., and Sjoholm, I., Detection and separation of lymphocytes with specific 

surface receptors by using microparticles, Biochem. J., 170, 161, 1978. 

14. Krause, H. J., Schwarz, A., and Rohdewald, P., Interfacial polymerization, a useful method for the 

preparation of polymethylcyanoacrylate nanoparticles, Drug Dev. Ind. Pharm., 12, 527, 1986. 

15. Radwan, M. A., Preparation and in vivo evaluation of parenteral metoclopramide-loaded poly(alkylcyanoacrylate) 

nanospheres in rats, J. Microencapsul., 18, 467, 2001. 

16. Couvreur, P. et al., Polycyanoacrylate nanocapsules as potential lysosomotropic carriers: Preparation, 

morphological and sorptive properties, J. Pharm. Pharmacol., 31, 331, 1979. 

17. Grislain, L. et al., Pharmacokinetics and biodistribution of a biodegradable drug-carrier, Int. 

J. Pharm., 15, 335, 1983. 

18. Behan, N., Birkinshaw, C., and Clarke, N., Poly n-butyl cyanoacrylate nanoparticles: a mechanistic 

study of polymerization and particle formation, Biomaterials, 22, 1335, 2001. 

19. Seijo, B. et al., Design of nanoparticles of less than 50 nm diameter: Preparation, characterization 

and drug loading, Int. J. Pharm., 62, 1, 1990. 

20. Reddy, L. H. and Murthy, R. S. R., Influence of polymerization technique and experimental variables 

on the particle properties and release kinetics of methotrexate from poly(butylcyanoacrylate) nanoparticles, 

Acta Pharm., 54, 103, 2004. 

21. Reddy, L. H. and Murthy, R. S. R., Study of influence of polymerization factors on formation of 

poly(butyl cyanoacrylate) nanoparticles and in vitro drug release kinetics, ARS Pharm., 45, 211, 

2004. 

22. Fallaouh, N. A. K. et al., Development of a new process for the manufacture of polyisobutylcyanoacrylate 

nanocapsules, Int. J. Pharm., 28, 125, 1986. 

23. Florence, A. T., Whateley, T. L., and Wood, D. A., Potentially biodegradable microcapsules with 

poly-2-cyanoacrylate membranes, J. Pharm. Pharmacol., 31, 422, 1979. 

24. Chouinard, F. et al., Preparation and purification of polyisohexylcyanoacrylate nanocapsules, Int. 

J. Pharm., 72, 211, 1991. 

25. Palumbo, M. et al., Improved antioxidant effect of idebenone-loaded polyethyl-2-cyanoacrylate 

nanocapsules tested on human fibroblasts, Pharm. Res., 19, 71, 2002. 

26. Kopf, H. et al., Studium der Mizellpolymerisation in Gegenwart niedermolekularer Arzneistoffe. I. 

Herstellung und Isolierung der Nanopartikel, Restmonomerenbestimmung, physikalischchemische 

Daten, Pharm. Ind., 38, 281, 1976. 

27. Laakso, T. and Stjarnkvistjoholm, I., Biodegradable microspheres. VI. Lysosomal release of covalently 

bound antiparasitic drugs from starch microparticles, J. Pharm. Sci., 76, 134, 1987. 

28. Beck, P. et al., Influence of polybutyl cyanoacrylate nanoparticles and liposomes on the efficacy and 

toxicity of the anticancer drug mitoxantrone in murine tumor models, J. Microencapsul., 10, 101, 

1993. 

29. Illum, L. et al., Evaluation of carrier capacity and release characteristic for poly(butyl-2-cyanoacrylate) 


30. Couvreur, P. et al., Adsorption of antineoplastic drugs to polyalkyl cyanoacrylate nanoparticles and 

their release in calf serum, J. Pharm. Sci., 68, 1521, 1979. 

31. Kreuter, J. and Hartman, H. R., Comparative study on the cytostatic effects and tissue distribution of 

5-fluorouracil in a free from and bound to polybutylcyanocrylate nanoparticles in sarcoma 180bearing 

mice, Oncology, 40, 363, 1983. 

32. Marchal-Heussler, I. et al., Antiglaucomatous activity of betaxolol chlorhydrate sorbed onto 

different isobutyl cyanoacrylate nanoparticle preparations, Int. J. Pharm., 58, 115, 1990. 

33. Brasseur, N., Brault, D., and Couvreur, P., Adsorption of hematoporphyrin onto polyalkyl cyanoacrylate 

nanoparticles: carrier capacity and drug release, Int. J. Pharm., 70, 129, 1991. 

34. Gaspar, R., Preat, V., and Roland, M., Nanoparticles of polyisohexyl cyanoacrylate (PIHCA) as 

carriers of primaquine: formulation, physico-chemical characterization and acute toxicity, Int. 

J. Pharm., 68, 111, 1991. 


282 


35. Verdun, C. et al., Development of a nanoparticle controlled-release formulation for human use, 


36. Harmia, T., Speiser, P., and Kreuter, J., Optimization of pilocarpoine loading onto nanoparticles by 

sorption procedures, Int. J. Pharm., 33, 45, 1986. 

37. Poupaert, J. H. and Couvreur, P. A., Computationally derived structural model of doxorubicin 

interacting with oligomeric polyalkylcyanoacrylate in nanoparticles, J. Control. Release, 92, 19, 

2003. 

38. Chauvierre, C. et al., Heparin coated poly(alkylcyanoacrylate) nanoparticles coupled to hemoglobin: 

A new oxygen carrier, Biomaterials, 25, 3081, 2004. 

39. Simeonova, M. et al., Study on the role of 5-fluorouracil in the polymerization of butylcyanoacrylate 

during the formation of nanoparticles, J. Drug Target., 12, 49, 2004. 

40. Stella, B. et al., Design of folic acid-conjugated nanoparticles for drug targeting, J. Pharm. Sci., 89, 

1452, 2000. 

41. Kreuter, J., Physicochemical characterization of polyacrylic nanoparticles, Int. J. Pharm., 14, 

43, 1983. 

42. Kreuter, J. et al., Polybutyl cyanoacrylate nanoparticles for the delivery of [ 75 SE]norcholesterol, Int. 

J. Pharm., 16, 105, 1983. 

43. Magenheim, B. and Benita, S., Nanoparticle characterization: A comprehensive physicochemical 

approach, STP Pharma Sci., 1, 121, 1991. 

44. Edstrom, R. D. et al., Viewing molecules with scanning tunneling and atomic force microscopy, 

FASEB J., 4, 3144, 1990. 

45. Zasadzinski, J. A. N., Scanning tunneling microscopy with applications to biological surfaces, 

Biotechniques, 7, 174, 1989. 

46. Fuchs, H. and Laschinski, R., Surface investigations with a combined scanning-electron scanning 

tunneling microscope, Scanning, 12, 126, 1990. 

47. Gedde, U. W., Thermal analysis of polymers, Drug Dev. Ind. Pharm., 16, 2465, 1990. 

48. Muller, R. H. et al., Alkyl cyanoacrylate drug carriers. I. Physicochemical characterization of 

nanoparticles with different chain length, Int. J. Pharm., 84, 1, 1992. 

49. Andreade, J. D. et al., Contact angles at the solid–water interface, J. Colloid Interface Sci., 72, 488, 

1979. 

50. Wilkins, D. J. and Myers, P. A., Studies on the relationship between the electrophoretic properties of 

colloids and their blood clearance and organ distribution in the rat, Br. J. Exp. Pathol., 47, 568, 1966. 

51. Baszkin, A. and Lyman, D. J., The interaction of plasma proteins with polymers. I. Relationship 

between polymer surface energy and protein adsorption/desorption, J. Biomed. Mater. Res., 14, 393, 

1980. 

52. Lindsay, R. M. et al., Blood surface interactions, Trans. Am. Soc. Artif. Intern. Organs, 26, 603, 

1980. 

53. Andreade, J. D. et al., Surface characterization of poly(hydroxyethyl methacrylate) and related 

polymers. I. Contact angle methods in water, J. Polym. Sci. Polym. Symp., 66, 131, 1979. 

54. El-Egakey, M. A., Bentele, V., and Kreuter, J., Molecular weights of polycyanoacrylate nanoparticles, 

Int. J. Pharm., 13, 349, 1983. 

55. Muller, R. H. et al., In vitro model for the degradation of alkylcyanoacrylate nanoparticles, Biomaterials, 

11, 590, 1990. 

56. Scherer, D., Robinson, J. R., and Kreuter, J., Influence of enzymes on the stability of polybutylcyanoacrylate 


57. Sommerfeld, P., Schroeder, U., and Sabel, B. A., Long-term stability of PBCA nanoparticle suspensions 

suggests clinical usefulness, Int. J. Pharm., 155, 201, 1997. 

58. Maincent, P. et al., Disposition kinetics and oral bioavailability of vincamine-loaded polyalkyl 

cyanoacrylate nanoparticles, J. Pharm. Sci., 75, 955, 1986. 

59. Kreuter, J., Peroral administration of nanoparticles, Adv. Drug Deliv. Rev., 7, 71, 1991. 

60. Damge, C. et al., New approach for oral administration of insulin with polyalkylcayanoacrylate 

nanocapsules as drug carrier, Diabetes, 37, 246, 1988. 

61. Damge, C. et al., Nanocapsules as carriers for oral peptide delivery, J. Control. Release, 13, 233, 

1990. 



62. Wade, C. and Leonard, F., Degradation of poly-(methyl-2-cyanoacrylates), J. Biomed. Mater. Res., 

6, 215, 1972. 

63. Lenaerts, V. et al., Degradation of poly(isobutylcyanoacrylate) nanoparticles, Biomaterials, 5, 65, 

1984. 

64. Douglas, S. J., Illum, L., and Davis, S. S., Particle size and size distribution of Poly(butyl-2-cyanoacrylate) 

nanoparticles. II. Influence of stabilizers, J. Colloid Interface Sci., 103, 154, 1985. 

65. Couvreur, P., Design of biodegradable polyalkylcyanoacrylate nanoparticles as a drug carrier, In 

Microspheres and Drug Therapy, Davis, S. S., Illum, L., McVie, J. G., and Tomlinson, E., Eds., 

Elsevier, Amsterdam, p. 103, 1984. 

66. Washington, C., Drug release from monodisperse systems: A critical review, Int. J. Pharm., 58, 1, 

1990. 

67. Leavy, M. V. and Benita, S., Drug release from submicron O/W emulsion: a new in vitro kinetic 

evaluation model, Int. J. Pharm., 66, 29, 1990. 

68. Calvo, P., Vila-Jato, J. L., and Alonso, M. J., Comparative in vitro evaluation of several colloidal 

systems, nanoparticles, nanocapsules and nanoemulsions as ocular drug carriers, J. Pharm. Sci., 85, 

530, 1996. 

69. El-Samaligly, M. S., Rohdewald, P., and Mohmoud, H. A., Polyalkyl cyanoacrylate nanocapsules, 

J. Pharm. Pharmacol., 38, 216, 1986. 

70. LeRay, A. M. et al., End chain radiolabeling and in vitro stability studies of radiolabeled poly 

(hydroxyl) nanoparticles, J. Pharm. Sci., 83, 845, 1994. 

71. Barratt, G., Colloid drug carriers: achievements and perspectives, Cell. Mol. Life Sci., 60, 21, 2003. 

72. Simeonova, M. et al., Tissue distribution of polybutylcyanoacrylate nanoparticles loaded with spinlabelled 

nitrosourea in Lewis lung carcinoma-bearing mice, Acta Physiol. Pharmacol. Bulg., 20, 77, 

1994. 

73. Reddy, L. H. and Murthy, R. S. R., Pharmacokinetics and bio-distribution of doxorubicin loaded 

poly(butyl) cyanoacrylate nanoparticles synthesized by two different techniques, Biomed. Pap., 148, 

1611, 2004. 

74. Reddy, L. H., Sharma, R. K., and Murthy, R. S.R, Enhanced tumor uptake of doxorubicin loaded 

poly(butyl cyanoacrylate) nanoparticles in mice bearing dalton’s lymphoma tumor, J. Drug Target., 

12, 443, 2004. 

75. Gipps, E. M. et al., Distribution of polyhexyl cyanoacrylate nanoparticles in nude mice bearing 

human osteosarcoma, J. Pharm. Sci., 75, 256, 1986. 

76. Illum, L. et al., Tissue distribution of poly(hexyl 2-cyanoacrylate) nanoparticles coated with monoclonal 

antibodies in mice bearing human tumor xenografts, J. Pharmacol. Exp. Ther., 230, 733, 

1984. 

77. Kante, B. et al., Tissue distribution of [ 3 H] actinomycin D adsorbed on polybutylcyanoacrylate 


78. Douglas, S. J., Davis, S. S., and Illum, L., Biodistribution of poly(butyl-2-cyanoacrylate) nanoparticles 

in rabbits, Int. J. Pharm., 34, 145, 1986. 

79. Lobenberg, R. et al., Body distribution of azidothymidine bound to hexylcyanoacrylate nanoparticles 

after i.v. injection to rats, J. Control. Release, 50, 21, 1998. 

80. Borchard, G. et al., Uptake of surfactant-coated poly(methylmethacrylate)-nanoparticles by bovine 

brain microvessel endothelial cell monolayers, Int. J. Pharm., 110, 29, 1994. 

81. Gulyaev, A. E. et al., Significant transport of doxorubicin into the brain with polysorbate 80-coated 

nanoparticles, Pharm. Res., 16, 1564, 1999. 

82. Woodcock, D. M. et al., Reversal of multidrug resistance by surfactants, Br. J. Cancer, 66, 62, 1992. 

83. Nerurkar, M. M., Burto, P. S., and Borchardt, R. T., The use of surfactants to enhance the permeability 

of peptides through Caco-2 cells by inhibition of an apically polarized efflux system, 

Pharm. Res., 13, 528, 1996. 

84. Ghanem, G. E. et al., Labelled polycyanoacrylate nanoparticles for human in vivo use, Appl. Radiat. 

Isot., 44, 1219, 1993. 

85. Skidan, I. N. et al., Modification of pharmacokinetics and antitumor efficacy of doxorubicin bound to 

polysorbate 80-coated polybutylcyanoacrylate nanoparticles. In Third Central European Conference 

on Human Tumor Markers, Biomarkers Environ. J., 4, 39, 2001. 


284 


86. Pappo, J. and Ermak, T. H., Uptake and translocation of fluorescent latex particles by rabbit Peyer’s 

patch follicle epithelium: a quantitative model for M cell uptake, Clin. Exp. Immunol., 76, 144, 1989. 

87. Eldridge, J. H. et al., Controlled release in the gut-associated lymphoid tissues. 1. Orally administered 

biodegradable microspheres target the Peyer’s patches, J. Control. Release, 11, 205, 1990. 

88. Dembri, A. et al., Targeting of 3 0 -Azido3 0 -deoxythymidine (AZT)-loaded poly(isohexylcyanoacrylate) 

nanospheres to the gastrointestinal mucosa and associated lymphoid tissues, Pharm. Res., 18, 

467, 2001. 

89. Reszka, R. et al., Body distribution of free, liposomal and nanoparticle-associated mitoxantrone in 

B16-melanoma-bearing mice, J. Pharmacol. Exp. Ther., 280, 232, 1997. 

90. Fattal, E. et al., Treatment of experimental salmonellosis in mice with ampicillin-bound nanoparticles, 

Antimicrob. Agents Chemother., 33, 1540, 1989. 

91. Chiannilkulchai, N. et al., Doxorubicin-loaded nanoparticles: increased efficacy in murine hepatic 

metastases, Selective Cancer Ther., 5, 1, 1989. 

92. Storm, G. et al., Surface modification of nanoparticles to oppose uptake by the mononuclear phagocyte 

system, Adv. Drug Deliv. Rev., 17, 31, 1995. 

93. Gref, R. et al., In Microparticulate Systems for the Delivery of Proteins and Vaccines, Chen, S. and 

Bernstein, H., Eds., Marcel Dekker, New York, p. 279, 1996. 

94. Gabizon, A., Selective tumor localization and improvement of therapeutic index of anthracyclines 

encapsulated in long circulating liposomes, Cancer Res., 52, 891, 1992. 

95. Gref, R., Biodegradable long-circulating polymeric nanospheres, Science, 263, 1600, 1994. 

96. Bazile, D. et al., Stealth MePEG-PLA nanoparticles avoid uptake by the mononuclear phagocyte 

system, J. Pharm. Sci., 84, 493, 1995. 

97. Peracchia, M. T. et al., Development of sterically stabilized poly(isobutyl 2-cyanoacrylate) nanoparticles 

by chemical coupling of poly(ethylene glycol), J. Biomed. Mater. Res., 34, 846, 1997. 

98. Peracchia, M. T. et al., Synthesis of novel MePEGcyanoacrylate–hexadecylcyanoacrylate amphiphilic 

copolymer for nanoparticle technology, Macromolecules, 30, 846, 1997. 

99. Peracchia, M. T. et al., Biodegradable PEGylated nanoparticles from a novel MePEGcyanoacrylate– 

hexadecylcyanoacrylate amphiphilic copolymer, Pharm. Res., 15, 548, 1998. 

100. Couvreur, P. et al., Tissue distribution of antitumor drugs associated with polyalkylcyanoacrylate 

nanoparticles, J. Pharm. Sci., 69, 199, 1980. 

101. Peracchia, M. T., Stealth w PEGylated polycyanoacrylate nanoparticles for intravenous administration 

and splenic targeting, J. Control. Release, 60, 121, 1999. 

102. Peracchia, M. T. et al., Visualization of in vitro protein rejecting properties of PEGylated stealth 

polycyanoacrylate nanoparticles, Biomaterials, 20, 1269, 1999. 

103. deGennes, P. G., Confirmation of polymers attached to an interface, Macromolecules, 13, 1069, 

1980. 

104. Peracchia, M. T. et al., Complement consumption by poly(ethylene glycol) in different confirmations 

chemically coupled to poly(isobutyl-2-cyanoacrylate) nanoparticles, Life Sci., 61, 749, 1997. 

105. Ping, L. Y. et al., PEGylated polycyanoacrylate nanoparticles as salvicine carriers: Synthesis, preparation, 

and in vitro characterization, Acta Pharmacol. Sin., 22, 645, 2001. 

106. Kante, B. et al., Toxicity of polalkylcyanoacrylate nanoparticles. I. Free nanoparticles, J. Pharm. 

Sci., 71, 790, 1982. 

107. Couvreur, P. et al., Toxicity of polyalkylcyanoacrylate nanoparticles. II. Doxorubicin-loaded nanoparticles, 

J. Pharm. Sci., 71, 786, 1982. 

108. Olivier, J. C. et al., Indirect evidence that drug brain targeting using polysorbate 80-coated polybutylcyanoacrylate 

nanoparticles is related to toxicity, Pharm. Res., 16, 1836, 1999. 

109. Simeonova, M., Chorbadjiev, K., and Antcheva, M., Study of the effect of polybutylcyanoacrylate 

nanoparticles and their metabolites on the primary immune response in mice to sheep red blood cells, 

Biomaterials, 19, 2187, 1998. 

110. Couvreur, P. et al., Biodegrdable polymeric nanoparticles as drug carrier for antitumor agents, In 

Polymeric Nanoparticles and Microspheres, Guiot, P. and Couvreur, P., Eds., CRC Press, Boca 

Raton, FL, pp. 27–93, 1986. 

111. Gelperina, S. E. et al., Toxicological studies of doxorubicin bound to polysorbate 80-coated poly 

(butyl cyanoacrylate) nanoparticles in healthy rats and rats with intracranial glioblastoma, Toxicol. 

Lett., 126, 131, 2002. 



112. Jefferiss, C. D., Lee, A. J., and Ling, R. S., Thermal aspects of self-curing polymethylmethacrylate, 

J. Bone Joint Surg. Br., 57, 511, 1975. 

113. Vinters, H. V. et al., The histotoxicity of cyanoacrylates: A selective review, Neuroradiology, 27, 

279, 1985. 

114. Vinters, H. V., Lundie, M. J., and Kaufmann, J. C., Long-term pathological follow-up of cerebral 

arteriovenous malformations treated by embolization with bucrylate, N. Engl. J. Med., 314, 477, 

1986. 

115. Tranbahuy, P. et al., Direct intratumoral embolization of juvenile angiofibroma, Am. J. Otolaryngol., 

15, 429, 1994. 

116. San Millan Ruiz, D. et al., Pathology findings with acrylic implants, Bone, 25(Suppl. 2), 85S, 

1999. 

117. Matsumoto, T., Carcinogenesis and cyanoacrylate adhesives, J. Am. Med. Assoc., 202, 1057, 1967. 

118. Samson, D. and Marshall, D., Carcinogenic potential of isobutyl-2-cyanoacrylate, J. Neurosurg., 65, 

571, 1986. 

119. Hatanaka, S. et al., Induction of malignant fibrous histiocytoma in female Fisher rats by implantation 

of cyanoacrylate, zirconia, polyvinyl chloride or silicone, In Vivo, 7, 111, 1993. 

120. Canter, H. I. et al., Tissue response to N-butyl-2-cyanoacrylate after percutaneous injection into 

cutaneous vascular lesions, Ann. Plast. Surg., 49, 520, 2002. 

121. Kattan, J. et al., Phase I clinical trial and pharmacokinetic evaluation of doxorubicin carried by 

polyisohexylcyanoacrylate nanoparticles, Invest. New Drugs, 10, 191, 1992. 

122. Gibaud, S. et al., Increased bone marrow toxicity of doxorubicin bound to nanoparticles, Eur. 

J. Cancer, 30, 820, 1994. 

123. Chiannilkulchai, N. et al., Hepatic tissue distribution of doxorubicin-loaded nanoparticles after i.v. 

administration in reticulosarcoma M5076 metastasis-bearing mice, Cancer Chemother. Pharmacol., 

26, 122, 1990. 

124. Zhang, Z. R., Liao, G. T., and Hou, S. X., Study on mitoxantrone polycyanoacrylate nanospheres, 

Yao Xue Xue Bao, 29, 544, 1994. 

125. Yang, Y. X. et al., Antitumor activity of mitoxantrone-nanosphere against murine liver tumor H22, 

Sichuan Da Xue Xue Bao Yi Xue Ban, 35, 68, 2004. 

126. Soma, C. E. et al., Investigation of the role of macrophages on the cytotoxicity of doxorubicin and 

doxorubicin-loaded nanoparticles on M5076 cells in vitro, J. Control. Release, 68, 283, 2000. 

127. Bogush, T. et al., Direct evaluation of intracellular accumulation of free and polymer-bound anthracyclines, 

Cancer Chemother. Pharmacol., 35, 501, 1995. 

128. Bennis, S. et al., Enhanced cytotoxicity of doxorubicin encapsulated in polyisohexylcyanoacrylate 

nanospheres against multidrug-resistant tumor cells in culture, Eur. J. Cancer, 30A, 89, 1994. 

129. Simeonova, M. and Antcheva, M., In vitro study of cytotoxic activity of vinblastine in a free form 

and associated with nanoparticles, Acta Physiol. Pharmacol. Bulg., 20, 31, 1994. 

130. Rusakov, I. G. et al., Use of cytostatics in polymer-drug complexes in the treatment of malignant 

neoplasms of the maxilla, Sov. Med., 1, 23, 1990. 

131. Pan, W. S. and Hu, J., Studies of 5-fluorouracil nanocapsules, Yao Xue Xue Bao, 26, 280, 1991. 

132. Jiang, X. H. et al., The antihepatoma effect of lyophilized aclacinomycin A polyisobutylcyanoacrylate 

nanoparticles in vitro and in vivo, Yao Xue Xue Bao, 30, 179, 1995. 

133. Brasseur, F. et al., Actinomycin D absorbed on polymethylcyanoacrylate nanoparticles: increased 

efficiency against an experimental tumor, Eur. J. Cancer, 16, 1441, 1980. 

134. Chissov, V. I. et al., Immediate results of the use of prospidin in a cyanoacrylate glue compound in 

cancer of the esophagus, Sov. Med., 7, 28, 1989. 

135. Pérez, M. et al., Use of N-butyl-2-cyanoacrylate in oral surgery: biological and clinical evaluation, 

Artif. Organs, 24, 241, 2000. 

136. Walker, J., Martin, C., and Callaghan, R., Inhibition of P-glycoprotein function by XR9576 in a solid 

tumor model can restore anticancer drug efficacy, Eur. J. Cancer, 40, 594, 2004. 

137. Cuvier, C. et al., Doxorubicin-loaded nanospheres bypass tumor cell multidrug resistance, Biochem. 

Pharmacol., 44, 509, 1992. 

138. Vauthier, C. et al., Poly(alkylcyanoacrylates) as biodegradable materials for biomedical applications, 



286 


139. Colin de Verdiere, A. et al., Uptake of doxorubicin from loaded nanoparticles in multidrug-resistant 

leukemic murine cells, Cancer Chemother. Pharmacol., 3, 504, 1994. 

140. Laurand, A. et al., Quantification of the expression of multidrug resistance-related genes in human 

tumor cell lines grown with free doxorubicin or doxorubicin encapsulated in polyisohexylcyanoacrylate 

nanospheres, Anticancer Res., 24, 3781, 2004. 

141. Pepin, X. et al., On the use of ion-pair chromatography to elucidate doxorubicin release mechanism 

from polyalkyl cyanoacrylate nanoparticles at the cellular level, J. Chromatogr. B, 702, 181, 1997. 

142. Hu, Y.-P. et al., On the mechanism of action of doxorubicin encapsulation in nanospheres for the 

reversal of multidrug resistance, Cancer Chemother. Pharmacol., 37, 556, 1996. 

143. Vauthier, C. et al., Drug delivery to resistant tumors: the potential of polyalkyl cyanoacrylate 

nanoparticles, J. Control. Release, 93, 151, 2003. 

144. Kobayashi, H., Takamura, Y., and Miyachi, H., Novel approaches to reversing anticancer drug 

resistance using gene-specific therapeutics, Hum. Cell., 14, 172, 2001. 

145. Juliano, R. L. et al., Antisense pharmacodynamics: critical issues in the transport and delivery of 

antisense oligonucleotides, Pharm. Res., 16, 494, 1999. 

146. Maia, R. C., Carrico, M. K., Klumb, C. E., Noronha, H., Coelho, A. M., Vasconcelos, F. C., and 

Ruimanek, V. M., Clinical approach to circumvention, of multidrug resistance in refractory 

leukaemic patients: Association of cyclosporine A with etoposide, J. Exp. Clin. Cancer Res., 16, 

419, 1997. 

147. Toffoli, G. et al., Reversal activity of cyclosporine A and its metabolites M1 M17 and M21 in 

multidrug resistant cells, Int. J. Cancer, 71, 900, 1997. 

148. Den Boer, M. L. et al., The modulating effect of PSC 833 cyclosporine A, verapamil and genistein on 

in vitro cytotoxicity and intracellular content of daunorubicin in childhood acute lymphoblastic 

leukemia, Leukemia, 12, 912, 1998. 

149. Soma, C. E. et al., Reversion of multidrug resistance by co-encapsulation of doxorubicin and 

cyclosporine A in polyalkyl cyanoacrylate nanoparticles, Biomaterials, 21, 1, 2000. 

150. Colin de Verdiere, A. et al., Reversion of multidrug resistance with polyalkyl cyanoacrylate nanoparticles; 

towards a mechanism of action, Br. J. Cancer, 76, 198, 1997. 

151. Soma, C. E. et al., Ability of doxorubicin-loaded nanoparticles to overcome multidrug resistance of 

tumor cells after their capture by macrophages, Pharm. Res., 16, 1710, 1999. 

152. Povarov, L. S. et al., Partial circumvention of P-glycoprotein (Pgp)-associated resistance to doxorubicin 

(Dox) in MCF7/R human breast carcinoma and P388/R murine leukemia cell lines by 

doxorubicin-14-O-hemiadipate (H-Dox), Russ. J. Bioorg. Chem., 21, 797, 1995. 

153. Leontieva, O. V., Preobrazhenskaya, M. N., and Bernacki, R. J., Partial circumvention of P-glycoprotein-mediated 

multidrug resistance by doxorubicin-14-O-hemiadipate, Invest. New Drugs, 20, 35, 

2002. 

154. Krishna, R. and Mayer, L. D., Liposomal doxorubicin circumvents PSC 833-free drug interactions, 

resulting in effective therapy of multidrug-resistant solid tumors, Cancer Res., 57, 5246, 1997. 

155. Goren, D. et al., Nuclear delivery of doxorubicin via folate-targeted liposomes with bypass of 

multidrug-resistance efflux pumps, Clin. Cancer Res., 6, 1949, 2000. 

156. Carr, B. I., Hepatocellular carcinoma: Current management and future trends, Gastroenterology, 

127, S218, 2004. 

157. Talmadge, J. E. and Hart, I. R., Inhibited growth of a reticulum cell sarcoma (M5076) Induced 

in vitro and in vivo by macrophage-activating agents, Cancer Res., 44, 2446, 1984. 

158. Talmadge, J. E. et al., Immunomodulatory and immunotherapeutic properties of recombinant 

g-interferon and recombinant tumor necrosis factor in mice, Cancer Res., 47, 2563, 1987. 

159. Iigo, M., Shimizu, I., and Sagawa, K., Synergistic antitumor effects of carboplatin and interferons on 

hepatic metastases of colon carcinoma 26 and M5076 reticulum cell sarcoma, Jpn. J. Cancer Res., 

84, 794, 1993. 

160. Barraud, L. et al., Increase of doxorubicin sensitivity by doxorubicin-loading into nanoparticles for 

hepatocellular carcinoma cells in vitro and in vivo, J. Hepatol., 42, 736, 2005. 

161. Zhi-Rong, Z. et al., Study on the anticarcinogenic effect and acute toxicity of liver-targeting mitoxantrone 

nanoparticles, World J. Gastroenterol., 5, 511, 1999. 

162. Tamai, I. and Tsuji, A., Drug delivery through the blood–brain barrier, Adv. Drug. Deliv. Rev., 19, 

401, 1996. 



163. Vajkoczy, P. and Menger, M. D., Vascular microenvironment in gliomas, J. Neurooncol., 50, 99, 

2000. 

164. Olivier, J.-C., Drug transport to brain with targeted nanoparticles. NeuroRx (The American Society 

for Experimental NeuroTherapeutics, Inc.), 2, 108, 2005. 

165. Rautioa, J. and Chikhale, P. J., Drug delivery systems for brain tumor therapy, Curr. Pharm. Des., 

10, 1341, 2004. 

166. Kreuter, J., Influence of the surface properties on nanoparticle-mediated transport of drugs to the 

brain, J. Nanosci. Nanotechnol., 4, 484, 2004. 

167. Calvo, P. et al., Quantification and localization of PEGylated polycyanoacrylate nanoparticles in 

brain and spinal cord during experimental allergic encephalomyelitis in the rat, Eur. J. Neurosci., 15, 

1317, 2002. 

168. Brigger, I. et al., Poly(ethylene glycol)-coated hexadecylcyanoacrylate nanospheres display a 

combined effect for brain tumor targeting, J. Pharmacol. Exp. Ther., 303, 928, 2002. 

169. Stouch, T. R. and Gudmundsson, O., Progress in understanding the structure–activity relationships of 

P-glycoprotein, Adv. Drug Deliv. Rev., 54, 315, 2002. 

170. Doyle, L. A. et al., A multidrug resistance transporter from human MCF-7 breast cancer cells, Proc. 


171. Aouali, N. et al., Enhanced cytotoxicity and nuclear accumulation of doxorubicin-loaded nanospheres 

in human breast cancer MCF7 cells expressing MRP1, Int. J. Oncol., 23, 1195, 2003. 

172. Henry-Toulme, N., Grouselle, M., and Ramaseilles, C., Multidrug resistance bypass in cells exposed 

to doxorubicin loaded nanospheres absence of endocytosis, Biochem. Pharmacol., 50, 1135, 1995. 

173. Marbeuf-Gueye, C. et al., Kinetics of anthracycline efflux from multidrug resistance protein-expressing 

cancer cell compared with P-glycoprotein-expressing cancer cells, Mol. Pharmacol., 53, 141, 

1998. 

174. Kisara, S. et al., Combined effects of buthionine sulfoximine and cepharanthine on cytotoxic activity 

of doxorubicin to multidrug-resistant cells, Oncol. Res., 7, 191, 1995. 

175. Nishioka, Y. and Yoshino, H., Lymphatic targeting with nanoparticulate system, Adv. Drug Deliv. 

Rev., 47, 55, 2001. 

176. Hashida, M., Egawa, M., Muranishi, S., and Sezaki, H., Role of intra-muscular administration of 

water-in-oil emulsions as a method for increasing the delivery of anticancer agents to regional 

lymphatics, J. Pharmacokinet. Biopharm., 5, 223, 1997. 

177. Parker, R. J., Priester, E. R., and Sieber, S. M., Comparison of lymphatic uptake, metabolism, 

excretion and biodistribution of free and liposome entrapped [ 14 C]cytosine b-D-arabinofuranoside 

following intraperitoneal administration to rats, Drug Metab. Dispos., 10, 40, 1982. 

178. Porter, C. J., Drug delivery to the lymphatic system, Crit. Rev. Ther. Drug Carrier Syst., 14, 333, 

1997. 

179. Reddy, L. H. et al., Influence of administration route on tumor uptake and biodistribution of etoposide 

loaded solid lipid nanoparticles in Dalton’s lymphoma tumor bearing mice, J. Control. Release, 

105, 185, 2005. 

180. Gupta, P. K. and Hung, C. T., Magnetically controlled targeted micro-carrier systems, Life Sci., 44, 

175, 1989. 

181. Widder, K. J. et al., Selective targeting of magnetic albumin microspheres containing low-dose 

doxorubicin: total remission in Yoshida sarcomabearing rats, Eur. J. Cancer Clin. Oncol., 19, 

135, 1983. 

182. Rudge, S. et al., Adsorption and desorption of chemotherapeutic drugs from a magnetically targeted 

carrier (MTC), J. Control. Release, 74, 335, 2001. 

183. Gao, H. et al., Preparation of magnetic polybutyl cyanoacrylate nanospheres encapsulated with 

aclacinomycin A and its effect on gastric tumor, World J. Gastroenterol., 10, 2010, 2004. 

184. Yeo, W. et al., A phase II study of doxorubicin (A) versus cisplatin (P)/interferon g-2b (I)/doxorubicin 

(A)/fluorouracil (F) combination chemotherapy (PIAF) for inoperable hepatocellular carcinoma 

(HCC), Proc. Am. Soc. Clin. Oncol., 23, 4026, 2004. 

185. Loewe, C. et al., Arterial embolization of unresectable hepatocellular carcinoma with use of cyanoacrylate 

and lipiodol, J. Vasc. Interv. Radiol., 13, 61, 2002. 


288 


186. Loewe, C. et al., Permanent transarterial embolization of neuroendocrine metastases of the liver 

using cyanoacrylate and lipiodol: Assessment of mid- and long-term results, Am. J. Roentgenol., 180, 

1379, 2003. 

187. Rand, T. et al., Arterial embolization of unresectable hepatocellular carcinoma with use of microspheres, 

lipiodol, and cyanoacrylate, Cardiovasc. Intervent. Radiol., 28, 313, 2005. 

188. Harman, M., Etlik, O., and Unal, O., Direct percutaneous embolization of a carotid body tumor with 

n-butyl cyanoacrylate: An alternative method to endovascular embolization, Acta Radiol., 45, 646, 

2004. 

189. Yamagami, T. et al., Value of transcatheter arterial embolization with coils and n-butyl cyanoacrylate 

for long-term hepatic arterial infusion chemotherapy, Radiology, 230, 792, 2004. 

190. Berghammer, P. et al., Arterial hepatic embolization of unresectable hepatocellular carcinoma using 

a cyanoacrylate/lipiodol mixture, Cardiovasc. Intervent. Radiol., 21, 214, 1998. 

191. Inaba, Y. et al., Right gastric artery embolization to prevent acute gastric mucosal lesions in patients 

undergoing repeat hepatic arterial infusion chemotherapy, J. Vasc. Interv. Radiol., 12, 957, 2001. 

192. Nadalini, V. F. et al., Therapeutic occlusion of the hypogastric arteries with isobutyl-2-cyanoacrylate 

in vesical and prostatic cancer, Radiol. Med. (Torino), 67, 61, 1981. 

193. Papo, J., Baratz, M., and Merimsky, E., Infarction of renal tumors using isobutyl-2 cyanoacrylate and 

lipiodol, AJR Am. J. Roentgenol., 137, 781, 1981. 

194. Carmignani, G., Belgrano, E., and Puppo, P., First clinical results with isobutyl-2-cyanoacrylate in 

the arterial embolisation of cancer of the kidney, Radiol. Med. (Torino), 64, 991, 1978. 

195. Abud, D. G. et al., Intratumoral injection of cyanoacrylate glue in head and neck paragangliomas, 

Am. J. Neuroradiol., 25, 1457, 2004. 

196. Derdeyn, C. P., Direct puncture embolization for paragangliomas: promising results but preliminary 

data, Am. J. Neuroradiol., 25, 1453, 2004. 

197. Patetsios, P. et al., Management of carotid body paragangliomas and review of a 30-year experience, 

Ann. Vasc. Surg., 16, 331, 2002. 

198. Muhm, M. et al., Diagnostic and therapeutic approaches to carotid body tumors. Review of 24 

patients, Arch. Surg., 132, 279, 1997. 

199. Marchesi, M. et al., Surgical treatment of paragangliomas of the neck, Int. Surg., 82, 394, 1997. 

200. Casasco, A. et al., Devascularization of craniofacial tumors by percutaneous tumor puncture, Am. 

J. Neuroradiol., 15, 1233, 1994. 

201. Pierot, L. et al., Embolization by direct puncture of hypervascularized oral tumors, Ann. Otolaryngol. 

Chir. Cervicofac., 111, 403, 1994. 

202. Schwab, G. et al., Antisense oligonucleotides adsorbed to polyalkylcyanoacrylate nanoparticles 

specifically inhibit mutated Ha-ras-mediated cell proliferation and tumorigenicity in nude mice, 

Proc. Natl. Acad. Sci. USA, 91, 10460, 1994. 

203. Lambert, G., Fattal, E., and Couvreur, P., Nanoparticulate systems for the delivery of antisense 

oligonucleotides, Adv. Drug Deliv. Rev., 47, 99, 2001. 

204. Chavany, C. et al., Polyalkylcyanoacrylate nanoparticles as polymeric carriers for antisense oligonucleotides, 

Pharm. Res., 9, 441, 1992. 

205. Zimmer, A., Antisense oligonucleotide delivery with poly-hexylcyanoacrylate nanoparticles as 

carriers, Methods: A Companion to Methods in Enzymol., 18, 286, 1999. 

206. Kircheis, R., Tumor targeting with surface-shielded ligand–polycation DNA complexes, J. Control. 

Release, 72, 165, 2001. 

207. Nakada, Y. et al., Pharmacokinetics and biodistribution of oligonucleotide adsorbed onto poly(isobutylcyanoacrylate) 

nanoparticles after intravenous administration in mice, Pharm. Res., 13, 38, 

1996. 

208. Lambert, G. et al., Effect of polyisobutylcyanoacrylate nanoparticles and lipofectin loaded with 

oligonucleotides on cell viability and PKC alpha neosynthesis in HepG2 cells, Biochimie, 80, 

969, 1998. 

209. Lambert, G. et al., Polyisobutylcyanoacrylate nanocapsules containing an aqueous core for the 

delivery of oligonucleotides, Int. J. Pharm., 214, 13, 2001. 

210. Schroeder, U. et al., Nanoparticle technology for delivery of drugs across the blood–brain barrier, 

J. Pharm. Sci., 87, 1305, 1998. 


16 

CONTENTS 

Aptamers and Cancer 

Nanotechnology 

Omid C. Farokhzad, Sangyong Jon, and Robert Langer 

16.1 Targeted Drug Delivery: An Introduction ........................................................................ 289 

16.2 Nucleic Acid Ligands (Aptamers): From Discovery to Practice ..................................... 291 

16.2.1 Physicochemical Properties of Aptamers ........................................................... 292 

16.2.2 Methods for Isolation of Aptamers..................................................................... 293 

16.2.3 Examples of Aptamers for Targeted Delivery ................................................... 294 

16.2.3.1 Alpha-v Beta-3 ................................................................................... 294 

16.2.3.2 Human Epidermal Growth Factor-3 .................................................. 295 

16.2.3.3 Prostate Specific Membrane Antigen ................................................ 295 

16.2.3.4 Nucleolin ............................................................................................ 295 

16.2.3.5 Sialyl Lewis X.................................................................................... 295 

16.2.3.6 Cytotoxic T-Cell Antigen-4 ............................................................... 296 

16.2.3.7 Fibrinogen-Like Domain of Tenascin-C............................................ 296 

16.2.3.8 Pigpen ................................................................................................. 296 

16.3 Aptamer–Nanoparticle Conjugates for Targeted Cancer Therapy................................... 296 

16.3.1 Properties of Nanoparticles for Conjugation to Aptamers ................................. 297 

16.3.1.1 Size of Nanoparticles ......................................................................... 297 

16.3.1.2 Polymers for Synthesis of Nanoparticles........................................... 298 

16.3.1.3 Charge of Nanoparticles..................................................................... 298 

16.3.1.4 Surface Modification of Nanoparticles .............................................. 299 

16.3.2 Conjugation Strategies for Nanoparticle–Aptamer Conjugates ......................... 300 

16.4 Aptamer–Drug Conjugates for Targeted Cancer Therapy ............................................... 301 

16.5 Aptamer–Nanoparticle Conjugates for Protein Detection................................................ 304 

16.5.1 Aptamer–QD Conjugates for the Detection of Proteins .................................... 304 

16.5.2 Aptamer–Gold-Nanoparticle Conjugates for the Detection of Proteins ............ 305 

16.6 Conclusion ......................................................................................................................... 306 


References..................................................................................................................................... 306 

16.1 TARGETED DRUG DELIVERY: AN INTRODUCTION 

With advances in nanotechnology, it is becoming increasingly possible to combine specialized 

delivery vehicles and targeting approaches to develop highly selective and effective cancer 


289

290 


therapeutic and diagnostic modalities. 1–4 In the case of therapeutic vehicles, drug targeting is 

expected to reduce the undesirable and sometimes life-threatening side effects common in anticancer 

therapy. Drug targeting can also enhance the efficacy of drugs because the drug is delivered 

directly to cancer cells, making it possible to deliver a cytotoxic dose of the drug in a 

controlled manner. 

Cancer targeting may be achieved passively by continuously concentrating drug encapsulated 

nanoparticles in the tumor interstitial space due to the enhanced permeability and retention (EPR) 

effect. 5,6 This process occurs because of a differentially large quantity of nanoparticles extravasting 

out of tumor microvasculature, leading to an accumulation of drugs in the tumor interstitium. 

Multiple factors influence the EPR, including active angiogenesis and high vascularity, defective 

vascular architecture, impaired lymphatic drainage, and extensive production of vascular 

mediators, such as bradykinin, nitric oxide, vascular endothelial growth factor (VEGF), prostaglandins, 

collagenase, and peroxynitrite. The ease of vehicle extravasation is a function of the 

maximum size of the transvascular transport pathways in tumor microvasculature and is determined 

mainly through the size of open interendothelial gap junctions and trans-endothelial channels. The 

pore cutoff size of these transport pathways has been estimated between 400 and 600 nm, and 

extravasation of liposomes into tumors in vivo suggests a cutoff size in the range of 400 nm. 8 As a 

general rule, particle extravasation is inversely proportional to size, and small particles (fewer than 

200 nm size) should be most effective for extravasating the tumor microvasculature. 8–10 Passive 

tumor targeting has several limitations, including the inability to achieve a sufficiently high level of 

drug concentration at the tumor site, resulting in lack of drug efficacy. 11 Additionally, the lack of 

targeting efficiency contributes to undesirable systemic adverse effects (for reviews, see). 4,12 

Active tumor targeting may be achieved by both local and systemic administration of specially 

designed vehicles that recognize biophysical characteristics that are unique to the cancer cells. 

Most commonly, this represents binding of vehicles to target antigens using ligands that recognize 

tumor-specific or tumor-associated antigens. Some of these therapeutic conjugates are now under 

clinical development or are in clinical practice today. In the case of local drug delivery, such as 

through the injection of delivery vehicles within an organ, it is possible to achieve a desired effect 

within a subset of cells, as opposed to a generalized effect on all the cells of the targeted organ. 

In the case of cancer, the cytotoxic effects of a therapeutic agent would be directed to cancer cells 

while minimizing harm to noncancerous cells within and outside of the targeted organ. For 

example, suicide-targeted gene delivery has been demonstrated to be effective in killing prostate 

cancer but not healthy muscle cells in xenograft mouse models of prostate cancer. 13 Similar 

approaches have been used to deliver docetaxel via intratumoral injection of targeted controlled 

release polymer vehicles to prostate cancer cells in nude mice. 14 This approach is particularly 

useful for primary tumors that have not yet metastasized, such as localized breast or 

prostate cancer. 

For metastatic cancers, the drug delivery vehicle would ideally be administered systemically 

because the location, abundance, and size of tumor metastasis within the body limits its visualization 

or accessibility, thus making local delivery approaches impractical. Despite the obvious 

advantages of systemic delivery, this approach is more challenging, and many physiological and 

biochemical barriers must be overcome for vehicles to reach the tumor and ultimately be capable of 

delivering therapeutically effective concentrations of the drugs directly to the cancer cells. One 

major challenge is the early systemic clearance of vehicles by the mononuclear phagocytic cells 

present in the liver, spleen, lung, and bone marrow. 15–18 This is in part due to the large percentage of 

cardiac output directed to these organs, and in part due to the dense population of macrophages and 

monocytes that home in these organs. Additionally, a variety of factors play an important role in the 

ultimate success of systemic targeted approaches, including those factors related to the 

biochemical and physical characteristics of the nanoparticles such as: (1) the chemical properties 

of the controlled release polymer system and the encapsulated drug; (2) the size of the 


Aptamers and Cancer Nanotechnology 291 

nanoparticles; (3) surface charge and surface hydrophilicity of nanoparticles; and (4) the chemical 

nature of the targeting molecules. 19 

Several classes of molecules have been utilized for targeting applications, including various 

forms of antibody-based molecules, 20 such as chimeric human–murine antibodies, humanized 

antibodies, single chain Fv generated from murine hybridoma or phage display, and minibodies. 58 

Multivalent antibody-based targeting structures, such as multivalent minibodies, single-chain 

dimers and dibodies, and multispecific binding proteins, including bispecific antibodies and 

antibody-based fusion proteins, have all been evaluated. More recently, other classes of ligands, 

such as carbohydrates 21 and nucleic acid ligands, also called aptamers, 22,23 have been used as escort 

molecules for targeted delivery applications. The concept of nucleic acid molecules acting as 

ligands was first described in the 1980s when it was shown that some viruses encode small 

structured RNA that binds to viral and host proteins with high affinity and specificity. These 

RNA nucleic acid ligands had evolved over time to enhance the survival and propagation of the 

viruses. Subsequently, it was shown that these naturally occurring RNA ligands can inhibit the viral 

replication and have therapeutic benefits. 24 Finally, methods were developed to perform in vitro 

evolution and to isolate novel nucleic acid ligands that bind to a myriad of important molecules for 

diverse applications in research and clinical practice. 25,26 

16.2 NUCLEIC ACID LIGANDS (APTAMERS): FROM 

DISCOVERY TO PRACTICE 

The feasibility of using antibodies for targeted therapy, particularly for oncologic diseases, has been 

demonstrated repeatedly in the literature. Rituximab (Rituxane) was the first therapeutic based on 

monocloncal antibodies to receive FDA approval in 1997 for treating patients with relapsed or 

refractory low-grade or follicular, CD20 positive, B-cell non-Hodgkin’s lymphoma. 27 Awide 

variety of antibody-based drugs are now under clinical development or are in clinical practice 

today. For example, denileukin diftitox (Ontake) is an FDA-approved immunotoxin for the 

treatment of cutaneous T-cell lymphoma. 28 Many other radioimmunoconjugates or chemoimmunoconjugates 

directed against cell surface antigens are currently in various stages of clinical and 

pre-clinical development. Despite the recent success of monoclonal antibodies as targeting 

moieties, the use of antibodies for drug targeting may have a number of potential disadvantages. 

First, antibodies are large molecules (w20 nm for intact antibodies) 29,30 and their use in developing 

nanoscale therapeutic and diagnostic tools may result in an increase in vehicle size without added 

advantage. Second, antibodies may be immunogenic. Despite the current engineering approaches to 

yield improved humanized antibodies, this problem is not universally solved. Third, the biological 

development of monoclonal antibodies can be difficult and unpredictable. For example, the target 

antigen may not be well tolerated by the animal used to produce the antibodies, or the target 

molecules may be inherently less immunogenic, making it difficult to raise antibodies against 

such targets (although, this problem is overcome with the use of phage display libraries). 31,32 

Fourth, the production of antibodies involves a biological process that can result in batchto-batch 

variability in their performance, particularly when production is scaled up. The ideal 

targeting molecule for the delivery of nanoscale therapeutic and diagnostic systems should, like 

monoclonal antibodies, bind with high affinity and specificity to a target antigen but overcome or 

ameliorate some of the problems associated with the use and production of monoclonal antibodies. 

Nucleic acid aptamers are a novel class of ligands 25,26 that are small, nonimmunogenic, easy to 

isolate, characterize and modify, and exhibit high specificity and affinity for their target antigen. 

Aptamers derive their name from the Latin word aptus, meaning “to fit.” In the short time since Jack 

Szostak and Larry Gold independently described the groundbreaking methodology for in vitro 

evolution of aptamers, this class of ligands has emerged as an important arsenal in research and 

clinical medicine. 22,33 Considering the many favorable characteristics of aptamers, which have 

resulted in their rapid progress into clinical practice, 33 researchers have begun to exploit this class 


292 

of molecules for targeted delivery of controlled release polymer drug delivery vehicles. Recently, 

we described the first proof-of-concept drug delivery vehicles utilizing aptamers for targeted 

delivery of drug encapsulated nanoparticles 23 and have gone on to the show efficacy of similarly 

designed nanoparticles against prostate cancer tumors in vivo (Farokhzad et al., submitted). 14 

16.2.1 PHYSICOCHEMICAL PROPERTIES OF APTAMERS 


Nucleic acid aptamers are single-stranded DNA, RNA, or unnatural oligonucleotides that fold into 

unique structures capable of binding to specific targets with high affinity and specificity. Aptamers 

are selected in vitro from a pool of (w10 14 –10 15 ) random oligonucleotides. 34 These molecules have 

a molecular weight in the 10–20 kD range, which is one order of magnitude lower than that of 

antibodies (150 kD). 35 The small size of aptamers (w5 nm for 30–60 base pair of aptamer) 36 is 

desirable when utilizing them as targeting molecules for the delivery of nanoscale drug delivery 

vehicles, as they do not substantially increase the overall size of the vehicle. When compared to 

small molecule ligands, aptamers have a larger surface area, offering more points of chemical 

contact with their targets. Aptamers fold through intramolecular interaction into tertiary conformations 

with specific binding pockets. These molecules bind with high specificity and affinity to a 

variety of target antigens with an equilibrium dissociation constant (K d) in the 10 pM–10 mM 

range. 37 Unlike antibodies, their synthesis and large scale production is an entirely chemical 

process that does not rely on biological systems. This can translate into lower manufacturing 

cost and decrease batch-to-batch variability during production, which is a significant advantage 

for the commercialization of this class of molecules. Furthermore, due to their small size, aptamers 

exhibit superior tissue penetration, 22 and because of their similarity to endogenous molecules, 

aptamers are believed to be less immunogenic than antibodies. 38 

Unlike anti-sense oligonucleotides, which are single-stranded nucleic acids that require cellular 

uptake and affect the synthesis of a targeted protein by hybridizing to the mRNAs that encodes it, 

aptamers may inhibit a protein’s function by directly binding to their targets, which may include 

other nucleic acids, proteins, peptides, and small molecules. Aptamers can be described by a 

sequence of approximately 15–60 nucleotides (A, U, T, C, and G). The conformation of the 

aptamer confers specificity for a target molecule by interacting with multiple domains or 

binding pockets. Small changes in the target molecule can foil interactions, and thus aptamers 

can distinguish between closely related but nonidentical targets. For example, specific RNAs were 

identified that have a high affinity for the bronchodilator theophylline (1,3-dimethylxanthine) yet 

exhibit a more than 10,000 times weaker binding affinity to caffeine (1,3,7-trimethylxanthine), 

which differs form theophylline only by the substitution of a methyl group at the nitrogen atom 

N7 position. 39 Based on their unique molecular recognition properties, aptamers have found great 

utility for applications in areas such as in vitro and in vivo diagnostics, analytical techniques, 

imaging, and therapeutics. 35,40,41 

Aptamers are highly stable and may tolerate a wide range of temperature, pH (w4–9), and 

organic solvents without loss of activity, but these molecules are susceptible to nuclease 

degradation and renal clearance in vivo. Therefore, their pharmacokinetic properties must be 

enhanced prior to in vivo applications. Several approaches have been adopted to optimize the 

properties of aptamers, such as: (1) capping their terminal ends; (2) substituting naturally occurring 

nucleotides with unnatural nucleotides that are poor substrates for nuclease degradation (i.e., 2 0 -F, 

2 0 -OCH3, or2 0 -NH2-modified nucleotides); (3) substituting naturally occurring nucleotides with 

hydrocarbon linkers; and (4) use of L-enantiomers of nucleotides to generate mirror image aptamers 

commonly referred to as spiegelmers. 42–45 Aptamers can also be stabilized using locked nucleic 

acid modifications to reduce conformational flexibility, 46 circularized, linked together in pairs, or 

clustered onto a substrate. Alternatively, a nuclease resistant aptamer may be selected de novo 

using a pool of oligonucleotides with 2 0 -F- or 2 0 -OCH 3-modified nucleotides. By combining some 

of these strategies, an aptamer’s half-life can be prolonged from several minutes to many hours. 35 



To prolong the rate of clearance of aptamers, their size may be increased by conjugation with 

polymers such as polyethylene glycol (PEG). 47 

16.2.2 METHODS FOR ISOLATION OF APTAMERS 

In vitro selection, 25 also called systematic evolution of ligands by exponential enrichment 

(SELEX), 26,48 is a protocol to isolate rare functional oligonucleotides (i.e., aptamers) from a 

pool of random oligonucleotides (Figure 16.1). Similar to phage display or other strategies used 

to isolate ligands from random libraries, SELEX is essentially an iterative selection and amplification 

protocol to isolate single-stranded nucleic acid ligands that bind to their target with high 

affinity and specificity. The complexity of the starting library is determined in part by the number of 

random nucleotides in the pool. For example, by using a library with 40 random nucleotides, a pool 

of 10 24 distinct nucleotides can be generated. Practically speaking, the number of ligands in the 

starting pool for in vitro selection is closer to 10 15 , representing 1 nmol of the library. 

In the initial step, a library of random nucleotides flanked by fixed nucleotides is generated by 

solid phase oligonucleotide synthesis. The random nucleotides serve to add complexity to the pool 

while the fixed sequences are utilized for polymerase chain reaction (PCR) amplification. In the 

(10 14 -10 15 oligonucleotides) 

Nucleic acid library 

Fixed Random Fixed 

Primer 1 

ACGGACTACGGTTGAG-40N-CGAATGCGAACGTACAGT 

Primer 2 

Repeat until maximum 

enrichment of ligands 

has been obtained 

Isolation of bound 

nucleotides 

Target molecule bound 

to a solid support 

Cloning of oligonucleotides 

and amplification 

of plasmids 

Isolation of individual 

aptamers and 

characterization 

Reiterative rounds 

of selection and 

amplification 

FIGURE 16.1 Schematic representation of SELEX. An oligonucleotide library is synthesized containing 

random sequences that are flanked by fixed sequences that facilitate PCR amplification. Target molecules 

are incubated with this pool of oligonucleotides, and bound and unbound oligonucleotides are partitioned. 

Bound oligonucleotides are isolated, and iterative rounds of selection and amplification are performed with 

increased stringency to isolate aptamers with high specificity and affinity for the target molecule. Oligonucleotide 

ligands representing the aptamers are subsequently cloned in plasmids, amplified, and sequenced. The net 

result of this enrichment process is a small number of highly specific aptamers that are isolated from a large 

library of random oligonucleotides. 


294 

case of isolating DNA aptamers, the oligonucleotide pool is incubated with the target of interest, 

and the bound fragments are partitioned and directly amplified by PCR system. The resulting pool is 

used in a follow-up round of selection and amplification, and the process is repeated until the 

affinity for the target antigen plateaus. Typically, this will be achieved in six to ten rounds of 

SELEX. After the last round of SELEX, aptamers are cloned in plasmids, amplified, sequenced, and 

their binding constants are determined (Figure 16.1). These aptamers may be subject to additional 

modification such as size minimization to truncate the nucleotides not necessary for binding 

characteristics and nuclease stabilization by replacing naturally occurring nucleotides with 

modified nucleotides (i.e., 2 0 -F pyrimidines, 2 0 -OCH3 nucleotides) that are poor substrates for 

endo- and exonuclease degradation. 

In contrast to the isolation of DNA aptamers, which require single-step amplification after 

portioning, the selection of RNA aptamers involves additional steps, including transcription of 

an RNA pool from the starting DNA library, reverse transcription of the partitioned RNA pool to 

generate a cDNA fragment and subsequent amplification of DNA, and transcription into RNA for 

the next round of selection. 34 The advantage of RNA SELEX, however, is that unnatural nucleotides 

such as 2 0 -F pyrimidines and 2 0 -OCH3 nucleotides may be used in the transcription of the 

RNA pool because these modified bases are utilized by RNA polymerase as substrate. Furthermore, 

mutant RNA polymerases have also been described that are capable of improved incorporation of 

modified bases during transcription. 49 The resulting modified RNA pool can be used for isolation of 

nuclease-stable RNA aptamers. Recently, a fully 2 0 -OCH 3-modified anti-VEGF aptamer was 

selected, and when conjugated to 40 kDa PEG, it demonstrated a circulating half-life of 23 h. 50 

Conversely, a DNA polymerase that can incorporate unnatural bases such as 2 0 -F and 2 0 -OCH3 has 

not been described, and, consequently, DNA aptamers must be nuclease stabilized after the 

SELEX procedure. 

16.2.3 EXAMPLES OF APTAMERS FOR TARGETED DELIVERY 

Since their original description in 1990, aptamers have been isolated to a wide variety of targets, 

including intracellular proteins, transmembrane proteins, soluble proteins, carbohydrates, and small 

molecule drugs. As of the submission of this chapter, 624 articles related to aptamers have been 

published, and more than 200 aptamers have been isolated (comprehensively listed in the Aptamer 

Database), 51 and one aptamer, macugen (pegaptanib sodium), against the VEGF165 protein was 

recently approved by the FDA in December, 2004 for the treatment of neovascular macular 

degeneration, 52,53 underscoring the rapid progress of aptamers from its original conception to 

clinical application. In choosing aptamers for targeting cancer cells, the aptamer must be directed 

towards receptors that are preferentially or exclusively expressed on the plasma membrane of 

cancer cells. Alternatively, they may be delivered to extracellular matrix molecules that are 

expressed preferentially in tumors. To date, several tumor-specific or tumor-associated aptamers 

have been isolated (reviewed by Pestourie et al.). 54 A partial listing of these aptamers, as well as 

their size and binding characteristics, are outlined below. 

16.2.3.1 Alpha-v Beta-3 


Alpha-v beta-3 (avb3) integrin is a transmembrane glycoprotein that mediates numerous processes, 

including cell migration, cell growth, tumor growth and metastasis, angiogenesis, and vascular 

healing. The expression of this protein on the endothelial cells of tumor neovasculature is dramatically 

increased, making avb3 an interesting protein for targeting approaches. 55 An 85-base-pair 

2 0 -flouropyrimidine RNA aptamer (Apt-avb3) has been shown to bind to the avb3 protein on the 

surface of endothelial cells in vitro and inhibit endothelial cell proliferation and survival. 56 

The future in vivo use of the Apt-avb3 may require post-SELEX optimization, including size 

minimization to facilitate large-scale synthesis, and enhance the pharmacokinetic properties of 



this molecule. This aptamer may be utilized for targeted cancer diagnostic and therapeutic 

applications. 

16.2.3.2 Human Epidermal Growth Factor-3 

Human epidermal growth factor-3 (HER-3) is a receptor tyrosine kinase that is over-expressed in 

several cancers. Over-expression of HER-3 is also associated with drug resistance in many HER-2 

over-expressed tumors, making HER-3 a candidate target for drug delivery. A30 is an RNA 

aptamer that was isolated against the extracellular domain of the HER-3 and is shown to inhibit 

heregluin dependent tyrosine phosphorylation of HER-2 and heregluin-induced growth response of 

MCF-7 cells at KiZ10 and 1 nM, respectively. 57 The future use of A30 for in vivo application may 

require post-SELEX optimization of this aptamer, including nuclease stabilization and 

size minimization. 

16.2.3.3 Prostate Specific Membrane Antigen 

Prostate specific membrane antigen (PSMA) encodes a folate carboxypeptidase. Its expression is 

tightly restricted to prostate acinar epithelium and is increased in prostatic intraepithelial neoplasia, 

prostatic adenocarcinoma, and in tumor-associated neovasculature. Two 2 0 -flouropyrimidine RNA 

aptamers, named xPSM-A9 and xPSM-A10, against the extracellular domain of the PSMA have 

been isolated and characterized. 58 The aptamer xPSM-A9 inhibits the enzymatic function of the 

PSMA noncompetitively with a K iZ2.1 nmol, and aptamer xPSM-A10 inhibits the enzymatic 

function of PSMA competitively with a KiZ11.9 nmol. Aptamer xPSM-A10 has also been truncated 

from 71 nucleotides to its current size of 56 nucleotides (w18kD). 

16.2.3.4 Nucleolin 

Nucleolin was originally described as a nuclear and cytoplasmic protein; however, a number of 

recent studies have shown that it can also be expressed at the cell surface. 59,60 Nucleolin is involved 

in the organization of the nuclear chromatin, rDNA transcription, packaging of the pre-RNA, 

ribosome assembly, nucleocytoplasmic transport, cytokinesis, nucleogenesis, and apoptosis. 

AS-1411 (formerly AGRO100) is an aptamer capable of making G-quadruplexes that bind to 

nucleolin on the cell surface 61 and interact with the nuclear factor kappa B (NFkB) essential 

modulator (NEMO) inside the cell. 62 The cytosolic localization of AS-1411 after binding to cell 

surface nucleolin may be exploited for the intracellular delivery of nanoparticles to cancer cells. 

The use of AS-1411 as a therapeutic modality has also shown promise for the treatment of cancer in 

humans, and Antisoma of United Kingdom is evaluating this aptamer in phase-I clinical trials. 63 

The therapeutic benefit of AS-1411 is presumably attributed to the disruption of the NFkB signaling 

inside the cells. 64 

16.2.3.5 Sialyl Lewis X 

Sialyl Lewis X (sLe x ) is a tetra saccharide glycoconjugate of transmembrane proteins that acts 

as a ligand for the selectin proteins during cell adhesion and inflammation. The sLe x is also 

abnormally overexpressed on the surface of cancer cells and may play a role in cancer cell 

metastasis. An RNA aptamer referred to as clone 5 has been isolated and shown to bind to the 

sLe x with sub-nanomolar affinity and block the sLe x /selectin mediated adhesion of HL60 

monocytic cell line in vitro. 65 The future use of clone 5 for in vivo targeted delivery applications 

may require post-SELEX optimization of this aptamer, including nuclease stabilization 

and size minimization. 


296 

16.2.3.6 Cytotoxic T-Cell Antigen-4 

Cytotoxic T-cell antigen-4 (CTLA-4) is a transmembrane protein that is expressed on the surface of 

activated T-cells and functions to attenuate the T-cell response by raising the threshold needed for 

T-cell activation. D60 (KdZ33 nM) is a size-minimized, 35-base-pair, 2 0 -fluoropyrimidine 

modified nuclease stable RNA aptamer that was selected against CTLA-4 and shown to inhibit 

CTLA-4 function in vitro and enhance tumor immunity in mice. 66 

16.2.3.7 Fibrinogen-Like Domain of Tenascin-C 

Tenascin-C is an extracellular matrix protein that is over-expressed during tissue remodeling 

processes, such as fetal development, and wound healing, as well as tumor growth. TTA1 is an 

aptamer (K dZ5 nM) that was isolated against the fibrinogen-like domain of Tenascin-C, 67 and may 

potentially be useful for cancer diagnostic and targeted therapeutic applications. 

16.2.3.8 Pigpen 

Pigpen is an endothelial protein of the Ewing’s sarcoma family that parallels the transition from 

quiescent to angiogenic phenotypes in vitro. Using YPEN-1 endothelial cells and N9 micorglial 

cells, respectively, in a selection and counter-selection strategy in SELEX, a DNA aptamer named 

III.1 was isolated that preferentially bound to YPEN-1 cells. 68 The III.1 was also shown to selectively 

bind to the microvessels of experimental rat glioblastoma. The isolation and characterization 

of the III.1 target identified pigpen as the target antigen. The use of III.1 aptamer for targeting the 

microvasculature of tumors is a potentially powerful means of delivering drugs to the site of 

the cancer. 

16.3 APTAMER–NANOPARTICLE CONJUGATES FOR TARGETED 

CANCER THERAPY 


Virtually every branch of medicine has been dramatically impacted by controlled drug delivery 

strategies during the past four decades, 3,69–72 including cardiology, 73 ophthalmology, 74 endocrinology, 

75 oncology, 76 immunology, 77 and orthopedics. 78 Drugs can be released in a controlled 

manner from within a material through surface or bulk erosion of the material, diffusion of the drug 

from the interior of the material, or swelling followed by diffusion or triggered by the environment 

or other external events, 2 such as changes in pH, 79 light, 80 temperature, 81 or the presence of an 

analyte, such as glucose. 82 In general, controlled-release polymer systems deliver drugs in the 

optimum dosage for long periods, thus increasing the efficacy of the drug, maximizing patient 

compliance and enhancing the ability to use highly toxic, poorly soluble, or relatively 

unstable drugs. 

The conjugation of aptamers to drug encapsulated nanoparticles results in targeted delivery 

vehicles for therapeutic application. These may include delivery of small molecule drugs, protein 

based drugs, nucleic acid drugs (anti-sense oligoneucleotide, RNAi or gene therapy), and targeted 

delivery of agents for neutron capture therapy or photodynamic therapy. Aptamers may also 

be bound to imaging agents to facilitate diagnosis and identification of tumor metastases. For 

example, it may be useful to bind aptamers to optical imaging agents including fluorophores 68 

and quantum dots (nanocrystals) 83 or MRI imaging agents such as magnetic nanoparticles 84,85 for 

detection of small foci of cancer metastasis. Multiplex systems comprising drug laden nanoparticle 

aptamer conjugates, together with imaging agents, represent a prospective avenue to 

future research. 



16.3.1 PROPERTIES OF NANOPARTICLES FOR CONJUGATION TO APTAMERS 

Nanoparticles are a particularly attractive drug delivery vehicle for cancer therapeutics because 

they can be synthesized to recognize tumor-specific antigens and deliver drugs in a controlled 

manner. 4,23 The design of targeted drug delivery nanoparticles combines drug encapsulated 

materials, such as biodegradable polymers, with a targeting moiety (Figure 16.2). Ideally, biodegradable 

targeted nanoparticles should be designed with the following parameters: (1) small size 

(preferably between 10 and 200 nm); (2) high drug loading and entrapment efficiency; (3) low rate 

of aggregation; (4) slow rate of clearance from the bloodstream; and (5) optimized targeting to the 

desired tissue with minimized uptake by other tissues. The following sections will discuss the 

various parameters that must be considered for engineering of nanoparticles for targeted drug 

delivery applications, including the development of nanoparticle–aptamer bioconjugates. 

This will include discussion of nanoparticle biomaterial, size, charge, and surface modification 

schemes to achieve the desired design parameters. It is important to note that a detailed review is 

beyond the scope of this chapter, and the reader is referred to the following reviews for further 

information: see references. 86–90 

16.3.1.1 Size of Nanoparticles 

Biodegradable polymer 

Chemotherapeutic 

Surface functionality 

Spacer group 

Targeting moiety 

FIGURE 16.2 Schematic representation of targeted drug delivery vehicle composed of polymeric nanoparticles 

that are surface-modified with targeting agents. 

The biodistribution pattern of nanoparticles, active nanoparticle targeting to tumor antigens, and 

passive nanoparticle targeting by EPR 91 are all greatly effected by the size of the nanoparticle. 

Passive targeting of the nanoparticle occurs because microvasculatures of tumors are more “leaky,” 

thus permitting selective permeation of nanoparticles into the desired tumor tissue. This phenomenon 

has been exploited to target liposomes, therapeutic and diagnostic nanoparticles, and drugpolymer 

conjugates to cancer tissue (reviewed by Maeda). 91 Smaller particles (fewer than 150 nm) 

are better suited for permeating through the leaky microvasculature of the tumor cells, and their 

more pronounced surface curvature may also reduce interaction with the receptors on the surface of 

macrophages and subsequent clearance of the particles. 12 Biodistribution studies using liposomes 

have shown that although particles that are larger than 200 nm are largely taken up by the spleen, 

those less than 70 nm are also efficiently cleared by the liver. 92 Taken together, the optimal 

nanoparticle size should be experimentally determined for each formulation because the interplay 

of various parameters (polymer system, encapsulated drug, surface charge, surface modification) 

makes it difficult to extrapolate the ideal nanoparticle size from seemingly similar studies. In our 

laboratory, we have tuned the size of polymeric nanoparticles made either by emulsions or nanoprecipitation 

to study the best formulation for a specific application. Generally, one can make use of 

the interaction of different solvents used for making the emulsions or tune the size of the nanoparticle 

by adjusting solvent ratios and polymer concentrations in solution. Our biodistribution 

studies using various sizes of poly(lactic-co-glycolic acid) (PLGA)–PEG nanoparticle–aptamer 

bioconjugates, has suggested a linear relationship with regards to uptake by liver and spleen 

such that smaller particles (w80 nm) are modestly better at avoiding uptake by these organs 

(unpublished results). 


298 

16.3.1.2 Polymers for Synthesis of Nanoparticles 

Controlled release biodegradable nanoparticles for clinical applications can be made from a wide 

variety of polymers, including poly(lactic acid) (PLA), 93 poly(glycolic acid) (PGA), PLGA, 94 

poly(orthoesters), 95 poly(caprolactone), 96 poly(butyl cyanoacrylate), 97 polyanhydrides, 98 and 

poly-N-isopropylacrylamide. 99 Although many fabrication methods exist, polymeric nanoparticles 

are frequently made using an oil-in-water emulsion, 18,100 which involves dissolving a polymer and 

drug in an organic solvent, such as methylene chloride, ethyl acetate, or acetone. The organic phase 

is mixed with an aqueous phase by vortexing and sonicating and then evaporated, forcing the 

polymer to precipitate as nanoparticles in the aqueous phase. The particles are then recovered by 

centrifugation and lyophilization. 

One of the considerations with respect to the material used for drug delivery is its ability to 

encapsulate drugs as well as degrade over the appropriate times. This subject has been an active 

area of investigation by our group and other investigators in academic and industry laboratories for 

several decades. The result has been an increasing arsenal of polymers with distinct encapsulation 

and release characteristics for a myriad of research, industrial, and clinical applications. 101,102 PGA 

and PLA are common biocompatible polymers that are used for many biomedical applications. 

PGA is hydrophilic because it lacks a methyl group and is more susceptible to hydrolysis, making 

this polymer easily degradable. Alternatively, PLA is relatively more stable in the body. 103 

Through these unique properties, polymers such as PLGA have been derived that are made from 

both glycolic acid and lactic acid components. The ability to change the ratio of these two components 

of the polymer can then be used to dramatically alter the rate of degradation. Therefore, by 

choosing the desired polymer system for the synthesis of nanoparticles the rate of degradation and 

subsequent release of the molecule may be tuned for the intended application. 

16.3.1.3 Charge of Nanoparticles 


Nanoparticle charge has been shown to be important for regulating its pharmacokinetic properties. 

For example, it has been shown that anionic and cationic liposomes activate the complement system 

through distinct pathways, suggesting that particle charge may impact particle opsonization and 

phagocytosis. 104 Cationic charge on liposomes has also been shown to reduce their circulating halflife 

in blood and to affect their biodistribution between the tumor microvasculature and interstitium 

without impacting overall tumor uptake. 105 Nanoparticles could be synthesized with charged 

surfaces either by using charged polymers, such as poly-L-lysine, polyethylenimine (PEI), or 

polysaccharides, or through surface modification approaches. For example, the layer-by-layer 

deposition of ionic polymers have been used to change surface properties of nanoparticles, such 

as quantum dots, by depositing ionic polymers of interest on the charged nanoparticle surfaces. 106 

Furthermore, surface charge of nanoparticles has been shown to regulate their biodistribution. 107 

For example, increasing the charge of cationic pegylated liposomes decreases their accumulation in 

the spleen and blood while increasing their uptake by the liver and an increasing in the accumulation 

of liposomes in tumor vessels. 105 These experiments suggest that optimizing surface 

physicochemical properties of nanoparticles to better match the biochemical and physiological 

features of tumors may enhance the intratumoral delivery of nanoparticles for systemic 

therapeutic approaches. 

For conjugation of the negatively charged aptamers to nanoparticles, the surface charge of the 

nanoparticle may be important. For example, we believe that direct immobilization of aptamers on 

cationic nanoparticles made from PEI may result in formation of aptamer–PEI complex that render 

the aptamer ineffective as a targeting molecule (unpublished observation). Therefore, neutral 

polymers such as PLA, PLGA or those with a more negative charge, such as polyanhyrides, may 

be most suitable for conjugation to aptamers. We have used PLA–PEG block copolymers to 

generate aptamer–nanoparticles bioconjugates. 23,108 One approach that may facilitate the use of 



a wider array of biomaterials for aptamer targeted drug delivery is through methods of “masking” 

the surface charge of the particles. For example, the addition of neutrally charged hydrophilic layer 

of PEG on the surface of the nanoparticles may facilitate the use of positively charged materials for 

the synthesis of nanoparticles. These cationic nanoparticles are particularly useful for gene delivery 

applications and thus may enable efficient targeted gene delivery using aptamers. 

16.3.1.4 Surface Modification of Nanoparticles 

Nanoparticle surface modification may also be used to engineer their interaction with the 

surrounding tissue. These interactions could be positive (i.e., targeting molecules) or negative 

(i.e., nonadhesive coatings). The surface modification of nanoparticles is particularly important 

because intravenously applied nanoparticles may be captured by macrophages before ever reaching 

the target site. Therefore, surface modifying particles to render them invisible to macrophages is 

essential to making long-circulating nanoparticles. 18,109 The ability to control the biodistribution of 

nanoparticles is particularly important for drug carrying nanoparticles because the delivery of drugs 

to the normal tissues can lead to toxicity. 16,17 

Hydrophilic polymers such as PEG, 18,109 polysaccharides, 110,111 and small molecules 112 can be 

conjugated on the surface of nanoparticles to engineer particles with desirable biodistribution 

characteristics. For example, to enhance the rate of circulation within the blood and minimize 

uptake by nondesired cell types, nanoparticles may be coated with polymers such as PEG. 18,109 

Various molecular weights and types of PEG (linear or branched) have been used to coat nanoparticles. 

113 PEG coatings are also useful for minimizing nanoparticle aggregation and can be used 

to prevent clogging of small vasculature and improve size-based targeting. More recently, novel 

approaches aimed at conjugating small molecules on nanoparticles using high-throughput methods 

have yielded nanoparticle libraries that could be subsequently analyzed for targeted delivery. 112 

The use of similar high-throughput approaches has significant potential in optimizing nanoparticle 

properties for cancer therapy. 

Surface modification of nanoparticles can be achieved in a multistep approach by first generating 

nanoparticles and subsequently modifying the surface of particles to achieve the desired 

characteristics. Alternatively, amphiphilic polymers may be covalently linked prior to generating 

nanoparticles to simultaneously control the surface chemistry as well as encapsulate drugs and 

COOH 

PLA PLA 

PEG 

Drugs 

EDC/NHS NH2 Aptamer 

Drugs 

FIGURE 16.3 A schematic outlining a conjugation reaction between aptamers and polymer nanoparticles 

containing an encapsulated drug. Through incorporating a COOH-terminated, PEG-functionalized surface on 

the nanoparticle, NH2-modified aptamers can be easily conjugated using simple aqueous chemistry. 


300 

eliminate the need for subsequent chemical modifications once the particle has been synthesized. 

This method may provide a more stable coating and better nanoparticle protection in contact with 

blood. For example, PLA, poly(caprolactone), and poly(cyanoacrylate) polymers, have been chemically 

conjugated to PEG polymers. 18,114,115 We have synthesized nanoparticles from amphiphilic 

copolymers composed of lipophilic PLGA and hydrophilic PEG (Figure 16.3) polymers where the 

PEG migrates to the surface of the nanoparticles in the presence of an aqueous solution. 18 Asimilar 

approach has also been used to generate pegylated PLA nanoparticles using PLA–PEG block 

copolymers. 23,108 These particles may be used to extend the nanoparticle residence times in circulation 

and enhance accumulation in tumor tissue through “passive targeting” and EPR effect. 

In the case of engineering nanoparticles for active targeting, the polymer and its coating should 

have functional groups for the attachment of targeting moieties (which may be bound directly to the 

nanoparticle surface or though a spacer group). The targeting molecules can enhance the molecular 

interaction of the nanoparticles with a subset of cells or tissue. 

16.3.2 CONJUGATION STRATEGIES FOR NANOPARTICLE–APTAMER CONJUGATES 


Covalent conjugation of aptamers to substrates or drug delivery vehicles can be achieved most 

commonly through succinimidyl ester–amine chemistry that results in a stable amide linkage 43,43 

or through maleimide–thiol chemistry. 23,108,116–120 Potential noncovalent strategies include affinity 

interactions (i.e., streptavidin–biotin) and metal coordination (i.e., between a polyhistidine tag at the 

end of the aptamer and Ni C2 chelates with immobilized nitrilotriacetic acid on the surface of the 

polymer particles). These covalent and noncovalent strategies have been used to immobilize a wide 

range of biomolecules, including proteins, enzymes, peptides, and nucleic acids to delivery vehicles. 

We believe that covalently linked bioconjugates may result in enhanced stability in physiologic 

salt and pH while avoiding the unnecessary addition of biological components (i.e., streptavidin), 

thus minimizing immunologic reactions and potential toxicity. For covalent conjugation, the 

aptamer is typically modified to carry a terminal primary amine or thiol group that is in turn 

conjugated, respectively, to activated carboxylic acid N-hydroxysuccinimide (NHS) ester or maleimide 

functional groups present on the surface of drug delivery vehicles. These reactions are carried 

out under aqueous conditions with a product yield of 80–90%. 121 One potential difficulty with 

maleimide–thiol chemistry is the oxidation of the thiol group attached to aptamers during storage 

(formation of S–S bond between two thiol-modified aptamers), resulting in dimers of aptamers that 

are not able to participate in the conjugation reaction with the malimide group on particles. 

This problem can be partially alleviated by using a reducing agent, such as tris(2carboxyethyl)phosphine 

(TCEP), b-mercaptoethanol, or dithiothreitol (DTT), during the conjugation 

reaction. A potential advantage of using NHS–amine chemistry is that the unreacted 

carboxylic acid groups on the particle surface make the particle surface charge (zeta potential) 

slightly negative, thus reducing nonspecific interaction between the negatively charged aptamers 

and the negative particle surface. Recently, controlled-release nanoparticles generated from 

PLA–PEG block copolymer with a terminal carboxylic acid group attached to the PEG were 

conjugated with primary-amine-terminated aptamers. 23,103 In this case, the hydrophilic PEG 

group facilitated the presentation of the carboxylic acid on the particle surface for conversion to 

activated carboxylic acid NHS ester and conjugation to the primary-amine-modified aptamers. 

The conjugation of aptamers to nanoparticles can be qualitatively confirmed by fluorescent 

microscopy or flow cytometry through the use of fluorescent probes, such as fluorescein isothiocyanate 

(FITC), that are conjugated directly to the aptamers or indirectly to complementary 

oligonucleotides that hybridize to the aptamers. 23 Alternatively, analytical approaches, such as 

x-ray photoemission spectroscopy (XPS), may be used for characterization of the nanoparticle 

surface to confirm the extent of conjugation. The presence of a hydrocarbon spacer group 

between the nanoparticle surface and the aptamer should improve the probability of interaction 

between the aptamer and its target. Furthermore, a consistent density of the aptamer on the surface 



of nanoparticles can potentially be achieved by utilizing an excess molar amount of aptamer relative to 

the reactive group on the nanoparticle surface during the conjugation reactions. However, the optimal 

density of targeting molecule on nanoparticle surface may need to be experimentally determined. 122 

We have used the covalent conjugation approach to demonstrate a proof-of-concept for nanoparticle–aptamer 

bioconjugates that target the PSMA on the surface of prostate cancer cells and are 

taken up by cells that express the PSMA protein specifically and efficiently. 23 We have also shown 

using a microfluidic system that these nanoparticle–aptamer conjugates are capable of binding to 

their target cells under flow conditions suggesting their suitability for in vivo targeted drug-delivery 

applications. 108 Most recently, we have demonstrated the in vivo efficacy of docetaxel-encapsulated 

nanoparticle–aptamer conjugates using a xenograft prostate cancer nude mouse model. 14 

These approaches have paved the way for the use of aptamers for targeted delivery of drug 

encapsulated nanoparticles to a myriad of human cancers. 

16.4 APTAMER–DRUG CONJUGATES FOR TARGETED CANCER THERAPY 

The clinical utility of radioimmunoconjugates and chemoimmunoconjugates in cancer treatment is 

well documented with several examples of each in clinical practice or under development at this 

time. 123–125 More recently aptamer conjugates for cancer therapeutic and diagnostic applications, 

14,23,108,116,118–120,126 biosensors, 127–129 flow cytometry, 130 ELISAs, 131,132 and capillary 

electrophoresis for separation technology have been discribed. 40,133 As noted in Section 16.2, 

aptamers have distinct advantages over antibodies, and, for the purpose of targeted delivery of 

toxins, their greatest advantage lies in their relatively smaller size and remarkable stability. 

The smaller size of aptamers, as compared to antibodies, allows for a more efficient tumor 

penetration of aptamers, translating to a more efficient delivery of a cytotoxic payload to tumor 

cells. 119 Our research group is interested in the use of aptamer–toxin conjugates for therapeutic 

applications, 126 and we believe that aptamer–toxin conjugates may result in future novel 

therapeutic applications in oncology and other areas of medicine. 

For the preparation of the aptamer–drug conjugates, conventional conjugation strategies that 

have been used for the preparation of immunoconjugates may also be employed with some modifications. 

125 For targeted radiotherapy, beta-emitting radioisotopes, such as 131 I, 90 Y, or 67 Cu, or 

alpha-emitting radioisotopes such as 213 Bi or 211 At, may be used to develop radioaptamerconjugates. 

These radioisotopes may be linked directly to aptamers or immobilized through metal 

chelators attached to aptamers. For targeted chemotherapy, drug molecules (i.e., doxorubicin, 

calicheamicin, or auristatin) may be covalently coupled to one or more sites on the aptamer 

using simple chemistry (i.e., formation of amide, disulfide, or hydrazone bond). 134,135 The development 

of aptamer conjugates is facilitated by site-selective functionalization of aptamers at their 3 0 - 

and/or 5 0 -terminus, making it possible to reproducibly attach an exact number of drugs in a sitespecific 

manner. 23,108,136 Aptamers may also be functionalized within the oligonucleotide chain 

using modified nucleotides with desired functional groups; however, this latter approach may 

negatively impact the aptamer conformation and function. In the case of conjugation of toxins to 

antibodies, the reaction is most commonly not site-directed, and this nonspecific conjugation 

approach may negatively impact the binding characteristics of antibodies. 136 Furthermore, unlike 

antibodies, the desired linkers or functional groups on aptamers for conjugation on other applications 

can be easily introduced during the chemical synthesis of these molecules. We have 

schematically demonstrated several proposed conjugation strategies between aptamers and drugs 

(Figure 16.4). Typical functional groups modifiable at the 5 0 -or3 0 -ends of aptamers are mostly 

nucleophiles, including amine (–NH2), sulfhydryl (–SH), and hydrazide (–C(O)NHNH2). Appropriate 

functional groups on the drugs for immobilization on aptamers may be carboxylic acid (for amide bond 

formation with amine-modified aptamers), maleimide or unsaturated carbonyl (for thioether formation 

with sulfhydryl groups on aptamers), activated disulfide (for example, S-pyridyl disulfide) for disulfide 


302 

3' 

5' 

X 

3' 3' 

X Y 

5' 5' 

X X 

aptamer 

Site-specific conjugation 

3' 3' 3' PEG 3' 

5' 5' 5' 5' 

(a) (b) (c) (d) 

formation with sulfhydryl groups on aptamers, and ketone for hydrazone formation with hydrazide on 

aptamers. 136 

The covalent bonds formed through the aforementioned strategies are all cleavable in vivo, 

allowing for the active drug to be released for therapeutic efficacy. The choice of conjugation 

strategy may allow for a spatial control of this release and improved clinical efficacy. In addition, by 

altering the size of aptamer–drug complex, the pharmacokinetics properties of the conjugate may 

be controlled, allowing for temporal control of the drug circulation and clearance. For example, a longer 

circulation time and reduced renal clearance may be achieved by linking the aptamer–drug conjugates 

to PEG polymers to increase the size of the complex above the threshold for renal clearance. 137 

This conjugate form may result in improved drug delivery to cancers as a result of increased systemic 

circulation time. Intracellular versus extracellular drug release also have unique sets of design 

considerations. 138 In the case of the disulfide linkage approach, the aptamer–drug conjugate is expected 

to be stable in plasma during circulation with subsequent intracellular cleavage of disulfide bond and 

drug release due to high intracellular concentration of glutathione (GSH). 139 Hydrazone linkage also 

has favorable properties for conjugation of drugs to aptamers, allowing for a pH sensitive release of the 

drug in a relatively more acidic environment of the tumor tissues. 134,140,141 Therefore, disulfide and 

hydrazone linkages may be considered when developing aptamer conjugates for cancer therapeutics. 

We have recently developed a novel strategy for generating aptamer–drug conjugates by 

intercalating drugs between nucleic acid bases of the aptamer (Figure 16.5). 126 Using doxorubicin 

(dox) as a model drug and the A10 PSMA aptamer as model aptamer, we have developed a stable 

aptamer–dox conjugate through physical interactions. The resulting physical conjugate demonstrated 

the following characteristics: (1) high discrimination between cancer cells expressing 

target antigen and cells lacking target antigen; and (2) efficient delivery of anti-cancer drug to 

target cancer cells. 126 Specifically, we developed aptamer–dox conjugates and demonstrated the 

stability of this system in vitro. Using LNCaP prostate cancer cell lines that express the PSMA 

antigen and PC3 prostate cancer cell lines that do not express the PSMA protein, we demonstrated a 

differential uptake of the aptamer–dox conjugate by LNCaP cells, translating to a more efficient 

cellular toxicity and anti-cancer efficacy in vitro (Figure 16.6). These results suggest that physical 

conjugate strategies such as intercalation may serve as a novel approach for developing 

aptamer–drug therapeutics. 

linker 

X,Y= functional groups 

−NH 2 , −SH 

−C(O)NHNH 2, etc. 


Z Z 

Drugs 

Z = −COOH, Maleimide 

−C(O)R 

FIGURE 16.4 Available conjugation strategies between end-functionalized aptamers and drug molecules. 

Panels (a) through (d) demonstrate a series of single- and double-conjugation strategies. 



PSMA aptamer 

+ 

Doxorubicin 

intercalation 

Physical conjugate 

FIGURE 16.5 A schematic diagram outlining the formation of physical conjugate between aptamers and 

drugs capable of intercalating into base pairs of aptamers. Doxorubicin, an anti-cancer drug, can be intercalated 

into a PSMA aptamer to form a 1:1 complex. 

FIGURE 16.6 The confocal laser scanning microscopy images of LNCaP (a) and PC3 cells (b) after 

treatments of the aptamer–doxorubicin physical conjugate (1.5 mM) for 2 h. 


304 

In addition to small molecule and radioisotopes, bio therapeutics such as small interfering 

RNAs (siRNA) 142–144 and protein-based toxins may also be conjugated to aptamers. For 

example, siRNAs can be coupled to aptamers through disulfide bond, allowing the release of 

siRNAs in the active form after cytosolic uptake by the targeted cells. A conjugate of aptamer– 

siRNA may represent a novel class of therapeutics with widespread application in medicine. 116 

16.5 APTAMER–NANOPARTICLE CONJUGATES FOR PROTEIN DETECTION 

Quantum dots (QDs) are nanometer-sized semiconductor crystals with tunable fluorescence that 

can be developed as sensitive biosensors or probes for in vivo imaging and diagnostic applications. 

145–147 Unlike conventional small molecule fluorophores and biological fluorophores, 

such as the green fluorescence protein (GFP), QDs possess several favorable characteristics, 

including: (1) greatly improved photostability (lack of photobleaching) that enable real time 

imaging of biomarkers over an extended period of time; (2) a broad excitation wavelength and a 

narrow emission wavelength enabling excitation of multiple QDs simultaneously using a single 

laser wavelength; (3) bright fluorescence with high quantum yield; and (4) ability to functionalize 

the surface of QDs with various molecules ranging from small molecules to antibodies and aptamers 

for desired targeted applications. Taken together, these favorable characteristics have made 

QDs an emerging class of imaging probes with broad applications in medicine. 

Metallic nanoparticles, such as gold-nanoparticles (Gd-NPs), are capable of absorbing or scattering 

light and are now used in many biological applications, particularly visualization of cellular 

and tissue components by electron microscopy. 148,149 More recently, it has been reported that 

antibody–Gd-NP conjugates 150,151 can detect proteins 152 or viruses 153 in solution through surface 

plasmon resonance effect, which results in color change depending on the status of aggregation of 

the particles. This simple phenomenon together with the fact that Gd-NPs are biocompatible makes 

Gd-NPs useful detection probes of various in vivo applications. 154 

The surface functionalization of QDs and Gd-NPs with specific targeting molecules may result 

in the development of targeted imaging modalities with widespread applications in research 

and medicine. 

16.5.1 APTAMER–QD CONJUGATES FOR THE DETECTION OF PROTEINS 

A QD–aptamer beacon for the detection of thrombin was recently reported, 155 and the principle of 

this detection system is schematically illustrated (Figure 16.7). Briefly, QDs were surface 

Q 

Q 

QD 

Q 

Q 

thrombin 

Quenched fluorescence Enhanced fluorescence 

QD 


+ 

Q 

Quencher−oligoDNA 

FIGURE 16.7 A schematic diagram outlining QD–aptamer beacon system for the detection of thrombin. 

Target protein replaces the bound quencher–oligoDNA by forming stable complex with aptamers on a QD, 

resulting in recovery of a native QD’s fluorescence. 



functionalized with an aptamer known to bind to the thrombin protein to develop a thrombindetecting 

QD–aptamer beacon. A second oligonucleotide, having complementary sequences to the 

thrombin aptamer and bearing a quencher, was synthesized and utilized to quench the system in the 

absence of thrombin. When the quencher oligonucleotide was incubated with the QD–aptamer 

conjugate in the absence of thrombin, the hybridization of the quencher oligonucleotide and the 

thrombin aptamer proximated the quencher molecules to the surface of the QD, resulting in 

quenching effect through fluorescence resonance energy transfer (FRET). 156,157 In the presence 

of thrombin (1 mM), the quencher oligonucleotide is displaced, resulting in restoration of the 

fluorescence of the QD–aptamer conjugate. Using this QD–aptamer beacon system, a 1 mM concentration 

of thrombin could be detected specifically. We postulate that multiplexed detection of a 

panel of protein targets may be performed in parallel by formation of QD–aptamer conjugates from 

different aptamers, and QDs possessing different emitting fluorescence profile. 158,159 In addition, 

considering a growing number of tumor-specific or tumor-associated aptamers that are available 

today, it may be increasingly possible to develop more specific and more efficient cancer diagnostic 

probes for a myriad of important cancers. We have recently developed a QD–aptamer conjugate 

using the A10 PSMA aptamer and have demonstrated the differential binding and uptake of these 

vehicles by cells that express the PSMA antigen (unpublished results). 

16.5.2 APTAMER–GOLD-NANOPARTICLE CONJUGATES FOR THE DETECTION OF PROTEINS 

A highly specific sensing system for the detection of platelet-derived growth factor (PDGF) has 

been developed using PDGF-specific, aptamer–conjugated gold-nanoparticles. 160 The principle of 

the detection using the conjugate is illustrated in Figure 16.8. The binding of soluble PDGF protein 

to Gd-NP–aptamer conjugates results in a change in the color of the solution due to the formation of 

aggregates between the Gd-NP conjugates. In the absence of PDGF, the conjugates are dispersed 

separately in solution (red), but in the presence of PDGF the Gd-NP aggregate closely to an interparticle 

distances less than the average diameter of each nanoparticle, causing a dipole–dipole 

coupling between particles as well as increased scattering, resulting in a measurable color 

change (purple). 161–163 Using this detection system, Gd-NP–aptamer conjugate could detect nanomolar 

concentrations of PDGF in solution. Because Gd-NPs are biocompatible, easy to develop, 

and easy to functionalize with aptamers, it is expected that bioconjugates utilizing aptamers and 

Gd-NPs as target molecular probes may result in novel cancer diagnostic modalities in the future. 

Gd-NP−Aptamer conjugates in dispersion 

target protein 

Gd-NP−Aptamer conjugates after aggregation 

FIGURE 16.8 A schematic representation of the aggregation of gold-nanoparticle–aptamer conjugates in the 

presence of target proteins. Aggregation causes color change from red to purple. 


306 

16.6 CONCLUSION 

Bioconjugates comprising nanoparticles and aptamers represent a potentially powerful tool for 

developing novel diagnostic and therapeutic modalities for cancer detection and treatment. As 

drug delivery vehicles for cancer therapy, nanoparticle–aptamer bioconjugates can be designed 

to target and be taken up by cancer cells for targeted delivery and controlled release of chemotherapeutic 

drugs over an extended time directly at the site of tumors. The successful achievement 

of this goal requires the isolation of aptamers that bind to the extracellular domain of antigens 

expressed exclusively or preferentially on the plasma membrane of cancer cells or on the extra 

cellular matrices of tumor tissue. In addition, nanoparticles would have to be designed with the 

optimized properties that facilitate targeting and delivery of the drugs to the desired tissues while 

avoiding uptake by the mononuclear phagocytic system in the body. 

The targeted delivery of chemotherapeutic drugs for cancer therapy may minimize their side 

effects and enhance their cytotoxicity to cancer cells, resulting in a better clinical outcome. We 

anticipate that the combination of controlled release technology and targeted approaches may 

represent a viable approach for achieving this goal. One major clinical advantage of targeted 

drug-encapsulated nanoparticle conjugates over drugs that are directly linked to a targeting 

moiety is that large amounts of chemotherapeutic drug may be delivered to cancer cells per each 

delivery and biorecognition event. Another advantage would be the ability to simultaneously 

deliver two or more chemotherapeutic drugs and release each in a predetermined manner, thus 

resulting in effective combination chemotherapy, which is common for the management of many 

cancers. Antibodies and peptides have been widely used for the targeted delivery of drug encapsulated 

nanoparticles; however, the translation of these vehicles into clinical practice has lagged 

behind our advances in the laboratory. This has been largely due to the immunogenicity and 

nonspecific uptake of nanoparticle–antibody bioconjugates by nontargeted cells, resulting in 

toxicity or poor efficacy of these conjugates. Nanoparticle–aptamer bioconjugates face some of 

the same challenges of nonspecific uptake after systemic administration and thus must be engineered 

with surface physicochemical characteristics to avoid toxicity to nontargeted cells. We 

believe that optimal particle size and surface properties to sufficiently decrease the rate of nonspecific 

particle uptake while achieving successful targeting must be determined experimentally on a 

case-by-case basis, as this also depends on the polymer system, the drug being encapsulated and the 

tumor microenvironment including its vascularity. However, the advantage of nanoparticle– 

aptamer bioconjugates over their antibody conjugate counterparts lies in the ease of aptamer 

isolation, lack of immunogenicity, and a decreased batch-to-batch variability attributable to the 

chemical nature of aptamer synthesis and production, which can facilitate the translation of optimally 

engineered nanoparticle–aptamer bioconjugates into clinical practice. 


The authors thank Jack Szostak, Benjamin Teply, Jeffrey Karp, Jianjun Cheng, Chris Cannizarro, 

Etgar Levy-Nissenbaum, and Andrej Luptak for helpful discussions and review of this manuscript. 

This work was supported by NIH/NCI CA119349, NIH/NIBIB EB003647, and by the David Koch 

Cancer Research Fund. 

REFERENCES 


1. Peppas, N. A., Intelligent therapeutics: biomimetic systems and nanotechnology in drug delivery, 


2. Langer, R. and Tirrell, D. A., Designing materials for biology and medicine, Nature, 428(6982), 

487–492, 2004. 

3. Langer, R., Drug delivery and targeting, Nature, 392(6679), 5–10, 1998. 

4. Ferrari, M., Cancer nanotechnology: opportunities and challenges, Nat. Rev. Cancer, 5(3), 161–171, 2005. 




chemotherapy: mechanism of tumoritropic accumulation of proteins and the antitumor agent 

smancs, Cancer Res., 46(12 Pt 1), 6387–6392, 1986. 



7. Maeda, H., The enhanced permeability and retention (EPR) effect in tumor vasculature: the key role 

of tumor-selective macromolecular drug targeting, Adv. Enzyme Regul., 41, 189–207, 2001. 

8. Yuan, F., Dellian, M., Fukumura, D., Leunig, M., Berk, D. A., Torchilin, V. P., and Jain, R. K., 

Vascular permeability in a human tumor xenograft: molecular size dependence and cutoff size, 

Cancer Res., 55(17), 3752–3756, 1995. 

9. Kong, G., Braun, R. D., and Dewhirst, M. W., Hyperthermia enables tumor-specific nanoparticle 

delivery: effect of particle size, Cancer Res., 60(16), 4440–4445, 2000. 



xenograft, Cancer Res., 54(13), 3352–3356, 1994. 

11. Brigger, I., Morizet, J., Laudani, L., Aubert, G., Appel, M., Velasco, V., Terrier-Lacombe, M. J. 

et al., Negative preclinical results with stealth nanospheres-encapsulated Doxorubicin in an orthotopic 

murine brain tumor model, J. Control Release, 100(1), 29–40, 2004. 

12. Brigger, I., Dubernet, C., and Couvreur, P., Nanoparticles in cancer therapy and diagnosis, Adv. Drug 

Deliv. Rev., 54(5), 631–651, 2002. 

13. Anderson, D. G., Peng, W., Akinc, A., Hossain, N., Kohn, A., Padera, R., Langer, R., and Sawicki, 

J. A., A polymer library approach to suicide gene therapy for cancer, Proc. Natl Acad. Sci. U.S.A., 

101(45), 16028–16033, 2004. 

14. Farokhzad, O. C., Cheng, J., Teply, B. A., Sherifi, I., Jon, S., Kantoff, P. W., Richie, J. P., and Langer, 

R., Targeted nanoparticle-aptamer bioconjugates for cancer chemotherapy in vivo, Proc. Natl Acad. 

Sci. U.S.A., 103(16), 6315–6320, 2006. 


Adv. Drug Del. Rev., 56(11), 1649–1659, 2004. 

16. Demoy, M., Gibaud, S., Andreux, J. P., Weingarten, C., Gouritin, B., and Couvreur, P., Splenic 

trapping of nanoparticles: complementary approaches for in situ studies, Pharm. Res., 14(4), 

463–468, 1997. 

17. Gibaud, S., Andreux, J. P., Weingarten, C., Renard, M., and Couvreur, P., Increased bone marrow 

toxicity of doxorubicin bound to nanoparticles, Eur. J. Cancer, 30A(6), 820–826, 1994. 

18. Gref, R., Minamitake, Y., Peracchia, M. T., Trubetskoy, V., Torchilin, V., and Langer, R., Biodegradable 

long-circulating polymeric nanospheres, Science, 263(5153), 1600–1603, 1994. 

19. Monsky, W. L., Fukumura, D., Gohongi, T., Ancukiewcz, M., Weich, H. A., Torchilin, V. P., Yuan, 

F., and Jain, R. K., Augmentation of transvascular transport of macromolecules and nanoparticles in 

tumors using vascular endothelial growth factor, Cancer Res., 59(16), 4129–4135, 1999. 

20. Guillemard, V. and Saragovi, H. U., Novel approaches for targeted cancer therapy, Curr. Cancer 

Drug Targets, 4(4), 313–326, 2004. 

21. Zhang, X. Q., Wang, X. L., Zhang, P. C., Liu, Z. L., Zhuo, R. X., Mao, H. Q., and Leong, K. W., 

Galactosylated ternary DNA/polyphosphoramidate nanoparticles mediate high gene transfection 

efficiency in hepatocytes, J. Control Release, 102(3), 749–763, 2005. 

22. Hicke, B. J. and Stephens, A. W., Escort aptamers: a delivery service for diagnosis and therapy, 

J. Clin. Invest., 106(8), 923–938, 2000. 

23. Farokhzad, O. C., Jon, S., Khademhosseini, A., Tran, T. N., Lavan, D. A., and Langer, R., Nanoparticle-aptamer 

bioconjugates: a new approach for targeting prostate cancer cells, Cancer Res., 

64(21), 7668–7672, 2004. 

24. Sullenger, B. A., Gallardo, H. F., Ungers, G. E., and Gilboa, E., Overexpression of TAR sequences 

renders cells resistant to human immunodeficiency virus replication, Cell, 63(3), 601–608, 1990. 

25. Ellington, A. D. and Szostak, J. W., In vitro selection of RNA molecules that bind specific ligands, 

Nature, 346(6287), 818–822, 1990. 

26. Tuerk, C. and Gold, L., Systematic evolution of ligands by exponential enrichment: RNA ligands to 

bacteriophage T4 DNA polymerase, Science, 249(4968), 505–510, 1990. 


308 


27. James, J. S. and Dubs, G., FDA approves new kind of lymphoma treatment. Food and drug administration, 

AIDS Treat. News,No 284 2–3, 1997. 

28. Foss, F. M., DAB(389)IL-2 (ONTAK): a novel fusion toxin therapy for lymphoma, Clin. Lymphoma, 

1(2), 110–116, discussion 117, 2000. 

29. Harris, L. J., Larson, S. B., Hasel, K. W., Day, J., Greenwood, A., and McPherson, A., The threedimensional 

structure of an intact monoclonal antibody for canine lymphoma, Nature, 360(6402), 

369–372, 1992. 

30. Harris, L. J., Skaletsky, E., and McPherson, A., Crystallographic structure of an intact IgG1 monoclonal 

antibody, J. Mol. Biol., 275(5), 861–872, 1998. 

31. Marks, J. D., Selection of internalizing antibodies for drug delivery, Methods Mol. Biol., 248, 

201–208, 2004. 

32. Marks, J. D., Ouwehand, W. H., Bye, J. M., Finnern, R., Gorick, B. D., Voak, D., Thorpe, S. J., 

Hughes-Jones, N. C., and Winter, G., Human antibody fragments specific for human blood group 

antigens from a phage display library, Biotechnology (N Y), 11(10), 1145–1149, 1993. 

33. Doggrell, S. A., Pegaptanib: the first antiangiogenic agent approved for neovascular macular 

degeneration, Expert Opin. Pharmacother., 6(8), 1421–1423, 2005. 

34. Fitzwater, T. and Polisky, B., A SELEX primer, Methods Enzymol., 267, 275–301, 1996. 

35. White, R. R., Sullenger, B. A., and Rusconi, C. P., Developing aptamers into therapeutics, J. Clin. 

Invest., 106(8), 929–934, 2000. 

36. Huang, D. B., Vu, D., Cassiday, L. A., Zimmerman, J. M., Maher, L. J., and Ghosh, G., Crystal 

structure of NF-kappaB (p50)2 complexed to a high-affinity RNA aptamer, Proc. Natl Acad. Sci. 

U.S.A., 100(16), 9268–9273, 2003. 

37. Hermann, T. and Patel, D. J., Adaptive recognition by nucleic acid aptamers, Science, 287(5454), 

820–825, 2000. 

38. Drolet, D. W., Nelson, J., Tucker, C. E., Zack, P. M., Nixon, K., Bolin, R., Judkins, M. B. et al., 

Pharmacokinetics and safety of an anti-vascular endothelial growth factor aptamer (NX1838) following 

injection into the vitreous humor of rhesus monkeys, Pharm. Res., 17(12), 1503–1510, 2000. 

39. Jenison, R. D., Gill, S. C., Pardi, A., and Polisky, B., High-resolution molecular discrimination by 

RNA, Science, 263(5152), 1425–1429, 1994. 

40. Tombelli, S., Minunni, M., and Mascini, M., Analytical applications of aptamers, Biosens. Bioelectron., 

20(12), 2424–2434, 2005. 

41. Charlton, J., Sennello, J., and Smith, D., In vivo imaging of inflammation using an aptamer inhibitor 

of human neutrophil elastase, Chem. Biol., 4(11), 809–816, 1997. 

42. Beigelman, L., McSwiggen, J. A., Draper, K. G., Gonzalez, C., Jensen, K., Karpeisky, A. M., Modak, 

A. S. et al., Chemical modification of hammerhead ribozymes. Catalytic activity and nuclease 

resistance, J. Biol. Chem., 270(43), 25702–25708, 1995. 

43. Aurup, H., Williams, D. M., and Eckstein, F., 2′-Fluoro- and 2′-amino-2′deoxynucleoside 

5′-triphosphates as substrates for T7 RNA polymerase, Biochemistry, 

31(40), 9636–9641, 1992. 

44. Pieken, W. A., Olsen, D. B., Benseler, F., Aurup, H., and Eckstein, F., Kinetic characterization of ribonuclease-resistant 

2′-modified hammerhead ribozymes, Science, 253(5017), 314–317, 1991. 

45. Eulberg, D. and Klussmann, S., Spiegelmers: biostable aptamers, Chem. biochem., 4(10), 979–983, 

2003. 

46. Schmidt, K. S., Borkowski, S., Kurreck, J., Stephens, A. W., Bald, R., Hecht, M., Friebe, M., 

Dinkelborg, L., and Erdmann, V. A., Application of locked nucleic acids to improve aptamer in 

vivo stability and targeting function, Nucleic Acids Res., 32(19), 5757–5765, 2004. 

47. Tucker, C. E., Chen, L. S., Judkins, M. B., Farmer, J. A., Gill, S. C., and Drolet, D. W., Detection and 

plasma pharmacokinetics of an anti-vascular endothelial growth factor oligonucleotide-aptamer 

(NX1838) in rhesus monkeys, J Chromatogr. B Biomed. Sci. Appl., 732(1), 203–212, 1999. 

48. Gold, L. and Tuerk, L., Methods of Producing Nucleic Acid Ligands, U.S.A., 1997. 

49. Chelliserrykattil, J. and Ellington, A. D., Evolution of a T7 RNA polymerase variant that transcribes 

2′-O-methyl RNA, Nat. Biotechnol., 22(9), 1155–1160, 2004. 

50. Burmeister, P. E., Lewis, S. D., Silva, R. F., Preiss, J. R., Horwitz, L. R., Pendergrast, P. S., 

McCauley, T. G. et al., Direct in vitro selection of a 2′-O-methyl aptamer to VEGF, 

Chem. Biol., 12(1), 25–33, 2005. 



51. Lee, J. F., Hesselberth, J. R., Meyers, L. A., and Ellington, A. D., Aptamer database, Nucleic Acids 

Res., 32, D95–D100, 2004. 

52. Gragoudas, E. S., Adamis, A. P., Cunningham, E. T., Feinsod, M., and Guyer, D. R., Pegaptanib for 

neovascular age-related macular degeneration, N Engl. J. Med., 351(27), 2805–2816, 2004. 

53. Vinores, S. A., Technology evaluation: pegaptanib. Eyetech/Pfizer, Curr. Opin. Mol. Ther., 5(6), 

673–679, 2003. 

54. Pestourie, C., Tavitian, B., and Duconge, F., Aptamers against extracellular targets for in vivo 

applications, Biochimie, 87(9-10), 921–930, 2005. 

55. McNeel, D. G., Eickhoff, J., Lee, F. T., King, D. M., Alberti, D., Thomas, J. P., Friedl, A. et al., Phase 

I trial of a monoclonal antibody specific for alphavbeta3 integrin (MEDI-522) in patients with 

advanced malignancies, including an assessment of effect on tumor perfusion, Clin. Cancer Res., 

11(21), 7851–7860, 2005. 

56. Mi, J., Zhang, X., Giangrande, P. H., McNamara, J. O., Nimjee, S. M., Sarraf-Yazdi, S., Sullenger, 

B. A., and Clary, B. M., Targeted inhibition of alphavbeta3 integrin with an RNA aptamer impairs 

endothelial cell growth and survival, Biochem. Biophys. Res. Commun., 338(2), 956–963, 2005. 

57. Chen, C. H., Chernis, G. A., Hoang, V. Q., and Landgraf, R., Inhibition of heregulin signaling by an 

aptamer that preferentially binds to the oligomeric form of human epidermal growth factor receptor- 

3, Proc. Natl Acad. Sci. U.S.A., 100(16), 9226–9231, 2003. 

58. Lupold, S. E., Hicke, B. J., Lin, Y., and Coffey, D. S., Identification and characterization of nucleasestabilized 

RNA molecules that bind human prostate cancer cells via the prostate-specific membrane 

antigen, Cancer Res., 62(14), 4029–4033, 2002. 

59. Deng, J. S., Ballou, B., and Hofmeister, J. K., Internalization of anti-nucleolin antibody into viable 

HEp-2 cells, Mol. Biol. Rep., 23(3-4), 191–195, 1996. 

60. Christian, S., Pilch, J., Akerman, M. E., Porkka, K., Laakkonen, P., and Ruoslahti, E., Nucleolin 

expressed at the cell surface is a marker of endothelial cells in angiogenic blood vessels, J. Cell Biol., 

163(4), 871–878, 2003. 

61. Dapic, V., Abdomerovic, V., Marrington, R., Peberdy, J., Rodger, A., Trent, J. O., and Bates, P. J., 

Biophysical and biological properties of quadruplex oligodeoxyribonucleotides, Nucleic Acids Res., 

31(8), 2097–2107, 2003. 

62. Barnhart, K. M., Laber, D. A., Bates, P. J., Trent, J. O., and Miller, D. M., In AGRO100: The translation 

from lab to clinic of a tumor-targeted nucleic acid aptamer. Proceedings from the American Society 

for Clinical Oncology Meeting, American Society for Clinical Oncology, New Orleans, LA, 2004. 

63. Laber, D. A., Sharma, V. R., Bhupalam, L., Taft, B., Hendler, F. J., and Barnhart, K. M. In Update on 

the First Phase I Study of AGRO100 in Advanced Cancer, American Society of Clinical Oncology 

Annual Meeting, Abstract 3064, 2005. 

64. Girvan, A. C., Barve, S. S., Thongboonkerd, V., Klein, J. B., and Pierce, W. M., In Inhibition of 

NFkB signaling by an oligonucleotide aptamer, Proceedings of the American Association of Cancer 

Research, 2004. 

65. Jeong, S., Eom, T., Kim, S., Lee, S., and Yu, J., In vitro selection of the RNA aptamer against the Sialyl 

Lewis X and its inhibition of the cell adhesion,Biochem. Biophys. Res. Commun., 281(1),237–243,2001. 

66. Santulli-Marotto, S., Nair, S. K., Rusconi, C., Sullenger, B., and Gilboa, E., Multivalent RNA 

aptamers that inhibit CTLA-4 and enhance tumor immunity, Cancer Res., 63(21), 7483–7489, 2003. 

67. Hicke, B. J., Marion, C., Chang, Y. F., Gould, T., Lynott, C. K., Parma, D., Schmidt, P. G., and 

Warren, S., Tenascin-C aptamers are generated using tumor cells and purified protein, J. Biol. Chem., 

276(52), 48644–48654, 2001. 

68. Blank, M., Weinschenk, T., Priemer, M., and Schluesener, H., Systematic evolution of a DNA 

aptamer binding to rat brain tumor microvessels. selective targeting of endothelial regulatory 

protein pigpen, J. Biol. Chem., 276(19), 16464–16468, 2001. 

69. Santini, J. T., Cima, M. J., and Langer, R., A controlled-release microchip, Nature, 397(6717), 

335–338, 1999. 

70. Edwards, D. A., Hanes, J., Caponetti, G., Hrkach, J., Ben-Jebria, A., Eskew, M. L., Mintzes, J., 

Deaver, D., Lotan, N., and Langer, R., Large porous particles for pulmonary drug delivery, Science, 

276(5320), 1868–1871, 1997. 

71. LaVan, D. A., McGuire, T., and Langer, R., Small-scale systems for in vivo drug delivery, Nature 

Biotechnol., 21(10), 1184–1191, 2003. 


310 


72. Elisseeff, J., McIntosh, W., Fu, K., Blunk, B. T., and Langer, R., Controlled-release of IGF-I and 

TGF-beta1 in a photopolymerizing hydrogel for cartilage tissue engineering, J. Orthop. Res., 19(6), 

1098–1104, 2001. 

73. Tanabe, K., Regar, E., Lee, C. H., Hoye, A., van der Giessen, W. J., and Serruys, P. W., Local drug 

delivery using coated stents: new developments and future perspectives, Curr. Pharm. Des., 10(4), 

357–367, 2004. 

74. Ebrahim, S., Peyman, G. A., and Lee, P. J., Applications of liposomes in ophthalmology, Surv. 

Ophthalmol., 50(2), 167–182, 2005. 

75. Haak, T., New developments in the treatment of type 1 diabetes mellitus, Exp. Clin. Endocrinol. 

Diabetes, 107(suppl. 3), S108–S113, 1999. 

76. Guerin, C., Olivi, A., Weingart, J. D., Lawson, H. C., and Brem, H., Recent advances in brain tumor 

therapy: local intracerebral drug delivery by polymers, Invest. New Drugs, 22(1), 27–37, 2004. 

77. Jiang, W., Gupta, R. K., Deshpande, M. C., and Schwendeman, S. P., Biodegradable poly(lactic-coglycolic 

acid) microparticles for injectable delivery of vaccine antigens, Adv. Drug Deliv. Rev., 

57(3), 391–410, 2005. 

78. Takahira, N., Itoman, M., Higashi, K., Uchiyama, K., Miyabe, M., and Naruse, K., Treatment 

outcome of two-stage revision total hip arthroplasty for infected hip arthroplasty using antibioticimpregnated 

cement spacer, J. Orthop. Sci., 8(1), 26–31, 2003. 

79. Little, S. R., Lynn, D. M., Ge, Q., Anderson, D. G., Puram, S. V., Chen, J., Eisen, H. N., and Langer, 

R., Poly-beta amino ester-containing microparticles enhance the activity of nonviral genetic 

vaccines, Proc. Natl Acad. Sci. U.S.A., 101(26), 9534–9539, 2004. 

80. Lendlein, A., Jiang, H., Junger, O., and Langer, R., Light-induced shape-memory polymers, Nature, 

434(7035), 879–882, 2005. 

81. Lendlein, A. and Langer, R., Biodegradable, elastic shape-memory polymers for potential biomedical 

applications, Science, 296(5573), 1673–1676, 2002. 

82. Zion, T. C. and Ying, J. Y., In Glucose responsive nanoparticle for crontrolled insulin delivery., 

Materials Research Society Fall Meeting, Boston, Dec 1–5, 2003. 

83. Chan, W. C. and Nie, S., Quantum dot bioconjugates for ultrasensitive nonisotopic detection, 

Science, 281(5385), 2016–2018, 1998. 

84. Harisinghani, M. G., Barentsz, J., Hahn, P. F., Deserno, W. M., Tabatabaei, S., van de Kaa, C. H., de 

la Rosette, J., and Weissleder, R., Noninvasive detection of clinically occult lymph-node metastases 

in prostate cancer, N. Engl. J. Med., 348(25), 2491–2499, 2003. 

85. Bonnemain, B., Superparamagnetic agents in magnetic resonance imaging: physicochemical characteristics 

and clinical applications. A review., J. Drug Target, 6(3), 167–174, 1998. 

86. Bala, I., Hariharan, S., and Kumar, M. N., PLGA nanoparticles in drug delivery: the state of the art, 


87. Drotleff, S., Lungwitz, U., Breunig, M., Dennis, A., Blunk, T., Tessmar, J., and Gopferich, A., 

Biomimetic polymers in pharmaceutical and biomedical sciences, Eur. J. Pharm. Biopharm., 

58(2), 385–407, 2004. 

88. Moghimi, S. M., Hunter, A. C., and Murray, J. C., Nanomedicine: current status and future prospects, 

Faseb J., 19(3), 311–330, 2005. 


tissue, Adv. Drug Deliv. Rev., 55(3), 329–347, 2003. 

90. Ravi Kumar, M., Hellermann, G., Lockey, R. F., and Mohapatra, S. S., Nanoparticle-mediated gene 

delivery: state of the art, Expert Opin. Biol. Ther., 4(8), 1213–1224, 2004. 


effect in macromolecular therapeutics: a review, J. Control Release, 65(1-2), 271–284, 2000. 

92. Liu, D., Mori, A., and Huang, L., Role of liposome size and RES blockade in controlling biodistribution 

and tumor uptake of GM1-containing liposomes, Biochim. Biophys. Acta, 1104(1), 95–101, 1992. 

93. Alonso, M. J., Gupta, R. K., Min, C., Siber, G. R., and Langer, R., Biodegradable microspheres as 

controlled-release tetanus toxoid delivery systems, Vaccine, 12(4), 299–306, 1994. 

94. Davda, J. and Labhasetwar, V., Characterization of nanoparticle uptake by endothelial cells, Int. 

J. Pharm., 233(1-2), 51–59, 2002. 

95. Deng, J. S., Li, L., Tian, Y., Ginsburg, E., Widman, M., and Myers, A., In vitro characterization of 

polyorthoester microparticles containing bupivacaine, Pharm. Dev. Technol., 8(1), 31–38, 2003. 



96. Molpeceres, J., Chacon, M., Guzman, M., Berges, L., and del Rosario Aberturas, M., A polycaprolactone 

nanoparticle formulation of cyclosporin-A improves the prediction of area under the curve 

using a limited sampling strategy, Int. J. Pharm., 187(1), 101–113, 1999. 

97. Sommerfeld, P., Sabel, B. A., and Schroeder, U., Long-term stability of PBCA nanoparticle suspensions, 

J. Microencapsul., 17(1), 69–79, 2000. 

98. Gao, J., Niklason, L., Zhao, X. M., and Langer, R., Surface modification of polyanhydride microspheres, 

J. Pharm. Sci., 87(2), 246–248, 1998. 

99. Huang, G., Gao, J., Hu, Z., St John, J. V., Ponder, B. C., and Moro, D., Controlled drug release from 

hydrogel nanoparticle networks, J. Control Release, 94(2-3), 303–311, 2004. 

100. Win, K. Y. and Feng, S. S., Effects of particle size and surface coating on cellular uptake of polymeric 

nanoparticles for oral delivery of anticancer drugs, Biomaterials, 26(15), 2713–2722, 2005. 

101. Langer, R., Drug delivery. Drugs on target, Science, 293(5527), 58–59, 2001. 

102. Langer, R. and Peppas, N. A., Advances in biomaterials, drug delivery, and bionanotechnology, 

AICHE J., 49(12), 2990–3006, 2003. 

103. Matsusue, Y., Hanafusa, S., Yamamuro, T., Shikinami, Y., and Ikada, Y., Tissue reaction of bioabsorbable 

ultra high strength poly (L-lactide) rod. A long-term study in rabbits, Clin. Orthop. Relat. 

Res., 317, 246–253, 1995. 

104. Chonn, A., Cullis, P. R., and Devine, D. V., The role of surface charge in the activation of the 

classical and alternative pathways of complement by liposomes, J. Immunol., 146(12), 4234–4241, 

1991. 

105. Campbell, R. B., Fukumura, D., Brown, E. B., Mazzola, L. M., Izumi, Y., Jain, R. K., Torchilin, 

V. P., and Munn, L. L., Cationic charge determines the distribution of liposomes between the 

vascular and extravascular compartments of tumors, Cancer Res., 62(23), 6831–6836, 2002. 

106. Jaffar, S., Nam, K. T., Khademhosseini, A., Xing, J., Langer, R., and Belcher, A. M., Layer-by-layer 

surface modification and patterned electrodeposition of quantum dots, Nano Letters, 4(8), 

1421–1425, 2004. 

107. Wang, Y., Ameer, G. A., Sheppard, B. J., and Langer, R., A tough biodegradable elastomer, Nat. 

Biotechnol., 20(6), 602–606, 2002. 

108. Farokhzad, O. C., Khademhosseini, A., Jon, S., Hermmann, A., Cheng, J., Chin, C., Kiselyuk, A., 

Teply, B., Eng, G., and Langer, R., Microfluidic system for studying the interaction of nanoparticles 

and microparticles with cells, Anal. Chem., 77(17), 5453–5459, 2005. 

109. Gref, R., Minamitake, Y., Peracchia, M. T., Domb, A., Trubetskoy, V., Torchilin, V., and Langer, R., 

Poly(ethylene glycol)-coated nanospheres: potential carriers for intravenous drug administration, 

Pharm. Biotechnol., 10, 167–198, 1997. 

110. Lemarchand, C., Gref, R., and Couvreur, P., Polysaccharide-decorated nanoparticles, Eur. J. Pharm. 

Biopharm., 58(2), 327–341, 2004. 

111. Lemarchand, C., Gref, R., Passirani, C., Garcion, E., Petri, B., Muller, R., Costantini, D., and 

Couvreur, P., Influence of polysaccharide coating on the interactions of nanoparticles with biological 

systems, Biomaterials, 27(1), 108–118, 2006. 

112. Weissleder, R., Kelly, K., Sun, E. Y., Shtatland, T., and Josephson, L., Cell-specific targeting of 

nanoparticles by multivalent attachment of small molecules, Nat. Biotechnol., 23(11), 1418–1423, 

2005. 

113. Mosqueira, V. C., Legrand, P., Morgat, J. L., Vert, M., Mysiakine, E., Gref, R., Devissaguet, J. P., 

and Barratt, G., Biodistribution of long-circulating PEG-grafted nanocapsules in mice: effects of 

PEG chain length and density, Pharm. Res., 18(10), 1411–1419, 2001. 

114. Lode, J., Fichtner, I., Kreuter, J., Berndt, A., Diederichs, J. E., and Reszka, R., Influence of surfacemodifying 

surfactants on the pharmacokinetic behavior of 14C-poly (methylmethacrylate) nanoparticles 

in experimental tumor models, Pharm. Res., 18(11), 1613–1619, 2001. 

115. Bazile, D., Prud’homme, C., Bassoullet, M. T., Marlard, M., Spenlehauer, G., and Veillard, M., 

Stealth Me.PEG-PLA nanoparticles avoid uptake by the mononuclear phagocytes system, J. Pharm. 

Sci., 84(4), 493–498, 1995. 

116. McNamara, J. O., Andrechek, E. R., Wang, Y., Viles, K. D., Rempel, R. E., Gilboa, E., Sullenger, 

B. A., and Giangrande, P. H., Cell type-specific delivery of siRNAs with aptamer-siRNA chimeras, 

Nat. Biotechnol., 24(8), 1005–1015, 2006. 


312 


117. Farokhzad, O. C., Karp, J. M., and Langer, R., Nanoparticle-aptamer bioconjugates for cancer 

targeting, Expert Opin. Drug Deliv., 3(3), 311–324, 2006. 

118. Chu, T. C., Shieh, F., Lavery, L. A., Levy, M., Richards-Kortum, R., Korgel, B. A., and Ellington, 

A. D., Labeling tumor cells with fluorescent nanocrystal-aptamer bioconjugates, Biosens. Bioelectron., 

21(10), 1859–1866, 2006. 

119. Chu, T. C., Marks, J. W., Lavery, L. A., Faulkner, S., Rosenblum, M. G., Ellington, A. D., and Levy, 

M., Aptamer:toxin conjugates that specifically target prostate tumor cells, Cancer Res., 66(12), 

5989–5992, 2006. 

120. Cheng, J., Teply, B. A., Sherifi, I., Sung, J., Luther, G., Gu, F. X., Levy-Nissenbaum, E., Radovic- 

Moreno, A. F., Langer, R., and Farokhzad, O.C., Formulation of functionalized PLGA-PEG nanoparticles 

for in vivo targeted drug delivery, Biomaterials, 2006. In press. 

121. Sehgal, D. and Vijay, I. K., A method for the high efficiency of water-soluble carbodiimide-mediated 

amidation, Anal. Biochem., 218(1), 87–91, 1994. 

122. Takae, S., Akiyama, Y., Otsuka, H., Nakamura, T., Nagasaki, Y., and Kataoka, K., Ligand Density 

Effect on Biorecognition by PEGylated Gold Nanoparticles: Regulated Interaction of RCA(120) 

Lectin with Lactose Installed to the Distal End of Tethered PEG Strands on Gold Surface, Biomacromolecules, 

6(2), 818–824, 2005. 

123. Polakis, P., Arming antibodies for cancer therapy, Curr. Opin. Pharmacol., 5(4), 382–387, 2005. 

124. Waldmann, T. A., Immunotherapy: past, present and future, Nat. Med., 9(3), 269–277, 2003. 

125. Wu, A. M. and Senter, P. D., Arming antibodies: prospects and challenges for immunoconjugates, 

Nat. Biotechnol., 23(9), 1137–1146, 2005. 

126. Bagalkot, V., Farokhzad, O. C., Langer, R., and Jon, S., Aptamer-Doxorubicin physical conjugate as 

a novel trageted drug delivery platform, Angew Chem. Int. Ed. Engl., 2006, In press. 

127. Nutiu, R. and Li, Y., Structure-switching signaling aptamers: transducing molecular recognition into 

fluorescence signaling, Chemistry, 10(8), 1868–1876, 2004. 

128. Srinivasan, J., Cload, S. T., Hamaguchi, N., Kurz, J., Keene, S., Kurz, M., Boomer, R. M. et al., ADPspecific 

sensors enable universal assay of protein kinase activity, Chem. Biol., 11(4), 499–508, 2004. 

129. Stojanovic, M. N. and Kolpashchikov, D. M., Modular aptameric sensors, J. Am. Chem. Soc., 

126(30), 9266–9270, 2004. 

130. Yang, X., Li, X., Prow, T. W., Reece, L. M., Bassett, S. E., Luxon, B. A., Herzog, N. K. et al., 

Immunofluorescence assay and flow-cytometry selection of bead-bound aptamers, Nucleic Acids 

Res., 31(10), e54, 2003. 

131. Baldrich, E., Restrepo, A., and O’Sullivan, C. K., Aptasensor development: elucidation of critical 

parameters for optimal aptamer performance, Anal. Chem., 76(23), 7053–7063, 2004. 

132. Yan, X. R., Gao, X. W., Yao, L. H., and Zhang, Z. Q., [Novel methods to detect cytokines by 

enzyme-linked oligonucleotide assay], Sheng Wu Gong Cheng Xue Bao, 20(5), 679–682, 2004. 

133. Schou, C. and Heegaard, N. H., Recent applications of affinity interactions in capillary electrophoresis, 

Electrophoresis, 27(1), 44–59, 2006. 

134. Dillman, R. O., Johnson, D. E., Shawler, D. L., and Koziol, J. A., Superiority of an acid-labile 

daunorubicin-monoclonal antibody immunoconjugate compared to free drug, Cancer Res., 48(21), 

6097–6102, 1988. 

135. Jaracz, S., Chen, J., Kuznetsova, L. V., and Ojima, I., Recent advances in tumor-targeting anticancer 

drug conjugates, Bioorg. Med. Chem., 13(17), 5043–5054, 2005. 

136. Hermanson, G. T., Bioconjugate Techniques, Academic Press, San Diego, 1996. 


theory to practice, Pharmacol. Rev., 53(2), 283–318, 2001. 

138. Payne, G., Progress in immunoconjugate cancer therapeutics, Cancer Cell, 3(3), 207–212, 2003. 

139. Saito, G., Swanson, J. A., and Lee, K. D., Drug delivery strategy utilizing conjugation via reversible 

disulfide linkages: role and site of cellular reducing activities, Adv. Drug Deliv. Rev., 55(2), 199–215, 2003. 

140. Trail, P. A., Willner, D., Bianchi, A. B., Henderson, A. J., TrailSmith, M. D., Girit, E., Lasch, S., 

Hellstrom, I., and Hellstrom, K. E., Enhanced antitumor activity of paclitaxel in combination with 

the anticarcinoma immunoconjugate BR96-doxorubicin, Clin. Cancer Res., 5(11), 3632–3638, 1999. 

141. Mosure, K. W., Henderson, A. J., Klunk, L. J., and Knipe, J. O., Disposition of conjugate-bound and 

free doxorubicin in tumor-bearing mice following administration of a BR96-doxorubicin immunoconjugate 

(BMS 182248), Cancer Chemother. Pharmacol., 40(3), 251–258, 1997. 



142. Lu, P. Y., Xie, F. Y., and Woodle, M. C., siRNA-mediated antitumorigenesis for drug target 

validation and therapeutics, Curr. Opin. Mol. Ther., 5(3), 225–234, 2003. 

143. Cheng, J. C., Moore, T. B., and Sakamoto, K. M., RNA interference and human disease, Mol. Genet. 

Metab., 80(1-2), 121–128, 2003. 

144. Cheng, J. C. and Sakamoto, K. M., The emerging role of RNA interference in the design of novel 

therapeutics in oncology, Cell Cycle, 3(11), 1398–1401, 2004. 

145. Smith, A. M., Gao, X., and Nie, S., Quantum dot nanocrystals for in vivo molecular and cellular 

imaging, Photochem. Photobiol., 80(3), 377–385, 2004. 

146. Gao, X., Cui, Y., Levenson, R. M., Chung, L. W., and Nie, S., In vivo cancer targeting and imaging 

with semiconductor quantum dots, Nat. Biotechnol., 22(8), 969–976, 2004. 

147. Gao, X., Yang, L., Petros, J. A., Marshall, F. F., Simons, J. W., and Nie, S., In vivo molecular and 

cellular imaging with quantum dots, Curr. Opin. Biotechnol., 16(1), 63–72, 2005. 

148. Hayat, M., Colloidal Gold: Principles, Methods and Applications, Academic Press, San Diego, 

1989. 

149. West, J. L. and Halas, N. J., Engineered nanomaterials for biophotonics applications: improving 

sensing, imaging, and therapeutics, Annu. Rev. Biomed. Eng., 5, 285–292, 2003. 

150. Loo, C., Hirsch, L., Lee, M. H., Chang, E., West, J., Halas, N., and Drezek, R., Gold nanoshell 

bioconjugates for molecular imaging in living cells, Opt. Lett., 30(9), 1012–1014, 2005. 


cancer imaging and therapy, Nano Lett., 5(4), 709–711, 2005. 

152. Xu, S., Ji, X., Xu, W., Li, X., Wang, L., Bai, Y., Zhao, B., and Ozaki, Y., Immunoassay using probelabelling 

immunogold nanoparticles with silver staining enhancement via surface-enhanced Raman 

scattering, Analyst, 129(1), 63–68, 2004. 

153. Driskell, J. D., Kwarta, K. M., Lipert, R. J., Porter, M. D., Neill, J. D., and Ridpath, J. F., Low-level 

detection of viral pathogens by a surface-enhanced Raman scattering based immunoassay, Anal. 

Chem., 77(19), 6147–6154, 2005. 

154. Copland, J. A., Eghtedari, M., Popov, V. L., Kotov, N., Mamedova, N., Motamedi, M., and 

Oraevsky, A. A., Bioconjugated gold nanoparticles as a molecular based contrast agent: implications 

for imaging of deep tumors using optoacoustic tomography, Mol. Imaging Biol., 6(5), 341–349, 

2004. 

155. Levy, M., Cater, S. F., and Ellington, A. D., Quantum-dot aptamer beacons for the detection of 

proteins, Chem. biochem., 6(12), 2163–2166, 2005. 

156. Meyer, T. and Teruel, M. N., Fluorescence imaging of signaling networks, Trends Cell Biol., 13(2), 

101–106, 2003. 

157. Mank, M., Reiff, D. F., Heim, N., Friedrich, M. W., Borst, A., and Griesbeck, O., A FRET-Based 

Calcium Biosensor with Fast Signal Kinetics and High Fluorescence Change, Biophys. J., 90(5), 

1790–1796, 2006. 

158. Chan, W. C., Maxwell, D. J., Gao, X., Bailey, R. E., Han, M., and Nie, S., Luminescent quantum dots 

for multiplexed biological detection and imaging, Curr. Opin. Biotechnol., 13(1), 40–46, 2002. 

159. Han, M., Gao, X., Su, J. Z., and Nie, S., Quantum-dot-tagged microbeads for multiplexed optical 

coding of biomolecules, Nat. Biotechnol., 19(7), 631–635, 2001. 

160. Huang, C. C., Huang, Y. F., Cao, Z., Tan, W., and Chang, H. T., Aptamer-modified gold nanoparticles 

for colorimetric determination of platelet-derived growth factors and their receptors, Anal. 

Chem., 77(17), 5735–5741, 2005. 

161. Elghanian, R., Storhoff, J. J., Mucic, R. C., Letsinger, R. L., and Mirkin, C. A., Selective colorimetric 

detection of polynucleotides based on the distance-dependent optical properties of gold nanoparticles, 

Science, 277(5329), 1078–1081, 1997. 

162. Storhoff, J. J., Marla, S. S., Bao, P., Hagenow, S., Mehta, H., Lucas, A., Garimella, V. et al., Gold 

nanoparticle-based detection of genomic DNA targets on microarrays using a novel optical detection 

system, Biosens. Bioelectron., 19(8), 875–883, 2004. 

163. Storhoff, J. J., Lucas, A. D., Garimella, V., Bao, Y. P., and Muller, U. R., Homogeneous detection of 

unamplified genomic DNA sequences based on colorimetric scatter of gold nanoparticle probes, Nat. 

Biotechnol., 22(7), 883–887, 2004. 


17 

CONTENTS 

Polymeric Micelles for 

Formulation of Anti-Cancer 

Drugs 

Helen Lee, Patrick Lim Soo, Jubo Liu, 

Mark Butler, and Christine Allen 

17.1 Introduction ........................................................................................................................318 

17.2 General Properties of Polymeric Micelles for Drug Formulation ....................................319 

17.2.1 Building Blocks for Polymeric Micelles .............................................................319 

17.2.2 Physico-Chemical Properties of Polymeric Micelles ..........................................321 

17.2.2.1 Size and Size Distribution ..................................................................321 

17.2.2.2 Morphology .........................................................................................321 

17.2.2.3 Stability ...............................................................................................322 

17.2.3 Preparation of Block Copolymer Micelle Formulations .....................................323 

17.2.4 Properties of Micelle Formulations .....................................................................324 

17.2.4.1 Drug Partitioning.................................................................................324 

17.2.4.2 Factors Affecting Drug Loading .........................................................324 

17.2.4.3 Factors Influencing Drug Release from Polymeric Micelles .............325 

17.3 Physical Characteristics of Polymeric Micelles for Clinical 

Applications in Cancer Therapy ........................................................................................326 

17.3.1 Doxorubicin ..........................................................................................................327 

17.3.1.1 NK911 .................................................................................................327 

17.3.1.2 SP1049C ..............................................................................................328 

17.3.2 Paclitaxel ..............................................................................................................329 

17.3.2.1 Genexol-PM ........................................................................................329 

17.3.2.2 NK105 .................................................................................................330 

17.4 Interactions of Polymeric Micelles with Cancer Cells......................................................330 

17.4.1 Drug Diffusion......................................................................................................331 

17.4.2 Cellular Internalization.........................................................................................332 

17.4.3 In Vitro Cytotoxicity ............................................................................................334 

17.4.4 Influence or Effect on Multi-Drug Resistance.....................................................335 

17.5 In Vivo Biological Performance of Block Copolymer Micelles ......................................336 

17.5.1 Fate of Polymeric Micelles In Vivo ....................................................................337 

17.5.2 In Vivo Drug Release Profiles .............................................................................338 

17.5.3 Toxicity, Pharmacokinetics, and Anti-Cancer Efficacy of Drugs 

Formulated in Block Copolymer Micelles ..........................................................339 

17.5.3.1 Doxorubicin Formulations (SP1049C and NK911) ...........................339 


317

318 

17.5.3.2 Paclitaxel Formulations (Genexol-PM and NK105) ..........................339 

17.6 Evolution of Polymeric Micelles .......................................................................................340 

17.6.1 Triggered Stimulation of Drug Activation or Release ........................................340 

17.6.1.1 Ultrasonic Irradiation ..........................................................................340 

17.6.1.2 Thermosensitive Systems ....................................................................341 

17.6.1.3 Photodynamic Therapy .......................................................................342 

17.6.2 Active Targeting...................................................................................................342 

17.6.2.1 Ligand Coupling..................................................................................343 

17.6.2.2 pH-Responsive Systems ......................................................................344 

17.7 Future Perspectives.............................................................................................................345 

Acknowledgments .........................................................................................................................345 

References .....................................................................................................................................345 



The use of conventional formulations for the systemic administration of many small molecule anticancer 

agents often results in an unsatisfactory therapeutic effect because of the narrow therapeutic 

window of these drugs. The low therapeutic index of these drugs is mostly attributed to the doselimiting 

toxicities that are associated with them and/or the excipients used in the formulation. 

Therefore, colloidal carriers that are composed of biocompatible and biodegradable materials 

with sustained drug release have been developed for formulation and delivery. Nanosized colloidal 

carriers have been shown to allow for increased accumulation of drug at tumors via the enhanced 

permeation and retention (EPR) effect. 1 As a result of the altered pharmacokinetics, the exposure of 

chemotherapeutic drugs to healthy tissues is reduced. Colloidal carriers including liposomes, 

nanoparticles, and polymeric micelles have been explored extensively for the delivery of a wide 

range of drugs as anti-cancer therapy. 1,2 

Polymeric micelles are nanosized assemblies of amphiphilic block copolymers that are suitable 

for the delivery of hydrophobic and amphiphilic agents. In an aqueous medium, micelles consist of 

a hydrophilic shell that minimizes clearance by the mononuclear phagocytic system (MPS) and a 

hydrophobic core that functions as a reservoir for hydrophobic drugs. In the past two decades, four 

polymeric micelle formulations loaded with chemotherapeutic drugs (i.e., NK911, SP1049C, 

Genexol-PM, and NK105) have entered clinical trial development. The results from the clinical 

studies have indicated that the polymeric micelle formulations reduce the toxicity associated with 

conventional formulations of these drugs that, in turn, results in a higher therapeutic index. Many 

preclinical fundamental studies have evaluated the relationships between the composition of the 

copolymers and the physico-chemical properties of the micelles. The properties of the micelles 

such as polymer–drug compatibility, thermodynamic and kinetic stability, and the drug release 

profiles have been shown to influence the in vivo performance and therapeutic effectiveness of the 

micelle-formulated drugs. These studies serve as guidelines for the optimization of polymeric 

micelles for clinical applications. 

This chapter is divided into six major sections and provides a discussion of the various aspects 

relating to use of polymeric micelles for formulation of anti-cancer drugs. Section 17.2 contains 

information on the physico-chemical properties of micelles (e.g., composition, size and size distribution, 

morphology, stability), micelle preparation techniques as well as drug loading and release 

properties of micelles. Section 17.3 highlights the preclinical development of several polymeric 

micelle formulations that are currently under clinical trial evaluation for use as cancer therapy. 

Specifically, the optimization of the physico-chemical characteristics of NK911, SP1049C, 

Genexol-PM, and NK105 are discussed. In Section 17.4, the interaction between polymeric 

micelles and cancer cells, including cellular internalization, in vitro cytotoxicity, and 


Polymeric Micelles for Formulation of Anti-Cancer Drugs 319 

chemosensitization of multi-drug resistant (MDR) cancer cells is reviewed. In Section 17.5, the 

physico-chemical properties of polymeric micelles are related to their in vivo performance such as 

in vivo drug release, pharmacokinetics, toxicity profiles, and anti-cancer efficacy. Finally, in 

Section 17.6, the development of several advanced polymeric micelle systems that are capable 

of targeted delivery to tumors via external stimuli (e.g., ultrasound, heat, light) or active targeting 

mechanisms (e.g., ligand coupling, pH-sensitive) are introduced. 

17.2 GENERAL PROPERTIES OF POLYMERIC MICELLES FOR DRUG 

FORMULATION 

In order to fully exploit the benefits of polymeric micelles for drug delivery, the drug-loaded 

micelle system requires extensive optimization in terms of its physico-chemical characteristics. 

Factors that may influence the performance of drug-loaded micelle systems include the selection of 

the building blocks to form the micelles, physico-chemical properties, drug partitioning, drug 

loading, and drug release profile. 

17.2.1 BUILDING BLOCKS FOR POLYMERIC MICELLES 

Amphiphilic block copolymers are the typical building blocks used for the formation of micelles. 

The term amphiphilic refers to a molecule that consists of both hydrophobic and hydrophilic 

segments. In aqueous solution, these block copolymers can self-assemble to form micelles 

consisting of a hydrophobic core and a hydrophilic corona or shell. 3–6 The hydrophilic corona 

(a) 

Solvent 

(b) 

Solvent 

(c) 

+ Water 

or buffer 

Remove 

Solvent 

+ Water 

or buffer 

+ 

Drug 

Remove 

Solvent 

+ Water 

or buffer 

or 

Hydrophilic corona 

Hydrophobic core 

FIGURE 17.1 Techniques for preparation of drug-loaded block copolymer micelles. (a) Copolymers and 

drugs are directly dissolved in water and form drug-loaded micelles spontaneously using the direct dissolution 

method. (b) In the evaporation method, copolymers and drugs are dissolved in an organic solvent. The solvent 

is evaporated, and drug-loaded micelles are formed with the addition of an aqueous solution. (c) In the dialysis 

method, copolymers and drugs are dissolved in an organic solvent that is miscible with water. Subsequent 

addition of water induces formation of micelles with swollen cores. Finally, the solvent is removed by 

evaporation or dialysis to form intact drug-loaded micelles. The structure of a drug-loaded micelle is also 

shown above. 


320 

TABLE 17.1 

Examples of Polymers That Are Commonly Used as Building Blocks for Block Copolymer 

Micelles in Drug Delivery 

Hydrophilic segment (corona-forming block) 

Poly(ethylene glycol) or 

poly(ethylene oxide) 

PEG or PEO CH2 CH2 O 

m 

Poly(N-vinyl-pyrrolidone) PVP CH2 CH 

N O 

m 

Hydrophobic segment (core-forming block) 

Poly(caprolactone) PCL 

Poly(D,L-lactide) PDLLA 

O 

C CH 

Poly(glycolide) PGA 

O 

C (CH2) 5 O 

n 

acts as the interface between the core and the exterior environment, and the hydrophobic core 

serves as the reservoir for the incorporation of lipophilic drugs (Figure 17.1). 

In drug delivery, biodegradable and biocompatible copolymers are chosen as a result of their 

degradation into non-toxic oligomers or monomers that are eventually absorbed in the body and/or 

eliminated. Many of the amphiphilic block copolymers used in drug delivery contain either a 

polyester, polyether, or a poly(amino acid) derivative as the hydrophobic block. A number of 

these hydrophobic polymers that have been used as the core of a micelle are listed in Table 17.1. 

Polyesters such as poly(D,L-lactide) (PDLLA), poly(glycolide) (PGA), and poly(3-caprolactone) 

(PCL) have been chosen as the core block because these polymers are biocompatible, biodegradable, 

and FDA approved for use in medical devices. Poly(propylene oxide) (PPO) is a polyether that 

has been used as the hydrophobic block in the well-studied Pluronic w family, the triblock 

CH 3 

O 

n 

O 

C CH2 O 

n 

O 

O 

Poly(D,L-lactide-co-glycolide) PLGA C CH2 O 

x 

C CH O 

y 

O 

O 

Poly(aspartic acid) PAsp NH C CH NH 

CH2 COOH 

x 

C CH2 CH NH 

COOH 

y 

Poly(b-benzyl-L-aspartate) PBLA NH 

O 

C CH NH 

n 

Pluronic CH2 COOC H2 w copolymer (triblock) 

O CH2 CH2 O CH CH2 n 

O CH 2 CH 2 

m/2 

m/2 

CH3 Poly(ethylene oxide)-b-poly(propylene oxide)-b-poly(ethylene oxide) 



CH 3


copolymers of poly(ethylene glycol)-block-poly(propylene oxide)-block-poly(ethylene glycol) 

(PEG-b-PPO-b-PEG). 7 Finally, poly(amino acid)s such as poly(aspartic acid) (PAsp) and poly 

(b-benzyl-L-aspartate) (PBLA) have also been widely explored as materials for the hydrophobic 

block because they can be easily modified synthetically. The wide range of materials available for 

selection as the hydrophobic block of the copolymer enables the preparation of micelles’ having a 

variety of microenvironments within their cores. 

In contrast, the hydrophilic block is important for stabilizing the micelle in its aqueous environment. 

The criteria for the selection of the hydrophilic block is that the polymer should be both 

uncharged and water soluble (e.g., poly(ethylene glycol), polyacrylamide, polyhydroxyethylmethacrylate, 

poly(N-vinyl-pyrrolidone) (PVP), and poly(vinyl alcohol)). 8 Poly(ethylene glycol) (PEG) 

is most commonly used as the hydrophilic segment in micelles. The molecular weights of PEG used 

in block copolymers typically range between 1000 and 12,000 g/mol with a chain length that is 

typically equal to or greater than the length of the core-forming block. 4 PEG can also be referred to 

as poly(ethylene oxide) (PEO), a polymer with a molecular weight that is greater than 20,000 g/ 

mol, whereas PEG refers to polymers with molecular weights below this value. 9 Because of its high 

water solubility and large excluded volume, PEG can stabilize the micelles by sterically excluding 

other polymers from the surface of the micelles. 1,2,10 In addition, PEG can inhibit the surface 

adsorption of biological components such as proteins, improve the residence time of the micelles 

in the circulation, and limit the micelle clearance by the MPS. 10–13 

The selection of building blocks for the micelle system includes the type of polymers and the 

length of the core- and corona-forming polymer blocks; these will influence the physico-chemical 

properties of the micelles and also affect the overall therapeutic efficacy of the delivery system. 

17.2.2 PHYSICO-CHEMICAL PROPERTIES OF POLYMERIC MICELLES 

In order to design an effective micellar drug delivery system, several key physico-chemical properties 

of the micelles should be considered as a means to optimize performance. These include 

micelle size and size distribution, morphology, and stability. 

17.2.2.1 Size and Size Distribution 

Block copolymer micelles typically range in size from 10 to 100 nm. As a result of their small size 

and the steric stabilization provided by the hydrophilic corona, micelles are less susceptible to MPS 

clearance. 14,15 In addition, their small size allows them to extravasate through the leaky endothelium 

and accumulate at the tumor interstitium. Therefore, micelles can passively target tumors 

via the enhanced permeability and retention (EPR) effect. 1,16 Micelles can also pass through the 

pores of the renal filtration system as individual polymer chains following disassembly. A renal 

threshold limit of less than 50,000 g/mol has been reported for polymeric carriers, 17 so it is 

important to design a micelle system where the individual polymer chains have a molecular 

weight that is lower than this threshold limit. The small size of the micelles also allows for 

sterilization by filtration in order to remove bacteria prior to intravenous administration. 

Filtration can also be used to reduce the size distribution of colloidal carriers that have broad 

distributions following initial preparation. Micelles tend to have a relatively narrow size distribution; 

however, they are prone to secondary aggregation because of insufficient coverage of the 

hydrophobic core by the hydrophilic chains. Dilution has been shown to be useful in breaking apart 

the secondary aggregates of primary micelles. 18 

17.2.2.2 Morphology 

When the corona-forming block of a copolymer is smaller than the core-forming block, crew-cut 

micelles are produced. 3 These copolymers may be used to prepare a wide range of morphologies 

including spheres, rods, lamellae, vesicles, large compound micelles, etc. 19,20 Of particular interest 


322 

are block copolymer vesicles that may prove to be effective for the simultaneous delivery of both 

hydrophilic and hydrophobic agents. 21,22 However, most of the block copolymer micelles used in 

drug delivery consist of spherical, star-type micelles as they are formed from copolymers that have 

a hydrophilic block that is of equal or greater length than the hydrophobic core-forming block. 3 A 

sample transmission electron micrograph of a population of spherical micelles is shown in 

Figure 17.2. 

17.2.2.3 Stability 

ES1 B, 100k.tif 

PEO–BTMC micelle 

PEO–BTMC micelle 

Print mag: 101000× @ 112 mm 

16:30 01/14/04 


100 nm 

HV=75 kV 

Direct mag: 100000× 

FIGURE 17.2 Transmission electron micrograph of block copolymer micelles formed from methoxy poly 

(ethylene glycol)-block-poly(5-benxyloxy-trimethylene carbonate) copolymers. (Reproduced with permission 

from Biomacromolecules 2004, 5, 1810–1817. Copyright 2004, American Chemical Society). 

The physical stability of block copolymer micelles is important because the early release of the 

drug into the bloodstream as a result of the disassembly of copolymer chains results in insufficient 

accumulation of the drug at the target site. It is critical that the copolymer micelles possess 

sufficient stability to remain intact in the circulation. The thermodynamic tendency for micelles 

to disassemble into individual chains is reflected by the critical micelle concentration (CMC) of the 

copolymer. Below the CMC, the copolymer is in the form of single polymer chains or unimers, and 

above the CMC, the copolymer exists as micelles in equilibrium with a small population of single 

chains. In this way, the CMC represents an important parameter as it determines the thermodynamic 

stability of the micelles during dilution. 23 Small molecule surfactants form micelles 

that tend to have higher CMC values that make them more susceptible to disassembly following 

dilution. In contrast, amphiphilic block copolymers tend to be more thermodynamically stable with 

lower CMC values in the range of 10 K7 –10 K6 M. 3,24,25 



The CMC has been found to be dependent on the nature and length of the hydrophobic block, 

the length of the hydrophilic block, and the total molecular weight of the copolymer. 26,27 Specifically, 

increasing the hydrophobicity of the core-forming block reduces the CMC and, in turn, 

improves the thermodynamic stability of the micelles. For example, Leroux et al. have demonstrated 

that increasing the hydrophobic PDLLA block from 27 to 55 mol% in a series of PVP-b- 

PDLLA-b-PVP copolymers results in a decrease in the CMC from 19.9 to 5.1 mg/L. 28 In contrast, 

when the hydrophilic block length is increased and the hydrophobic block is kept constant, the 

CMC increases as has been shown for the Pluronic w copolymers. 29 

Even at concentrations below the CMC, micelles can remain kinetically stable for extended 

periods of time depending on their glass transition temperature (Tg) and/or Tm if the core-forming 

block is semicrystalline. Micelles with a core-forming block that has a lower Tg (below 378C) will 

tend to disassemble at a faster rate than those with a higher Tg as a result of the free motion of the 

core-forming polymer chains at temperatures above the Tg. For example, Kataoka et al. found a 

dramatic increase in the CMC of PEG-b-PDLLA micelles at a temperature above the Tg of the 

hydrophobic PDLLA polymer that suggested a decrease in the kinetic stability of the system. 30 

Kang et al. improved the kinetic stability of PEG-b-poly(lactide) (PLA) block copolymer micelles 

by blending equimolar mixtures of PEG-b-poly(D-lactide) (PDLA) and PEG-b-poly(L-lactide) 

(PLLA) enantiomeric copolymers. They reported superior kinetic stability as a result of the 

increased van der Waals interactions between the PLA polymer chains that promoted a denser 

packing and a tighter conformation. 31 Finally, the presence of the hydrophobic drug in the core has 

also been shown to promote kinetic stability of the formulation. 32 

17.2.3 PREPARATION OF BLOCK COPOLYMER MICELLE FORMULATIONS 

The selection of the appropriate preparative technique for micelle formation is largely based on 

the solubility of the copolymer and the drug in aqueous media. There are several methods 

that have been developed for the preparation of drug-loaded block copolymer micelles, 

including the direct dissolution method, evaporation method, and dialysis method as depicted 

in Figure 17.1. 

The direct dissolution method is employed for block copolymers that are relatively soluble in 

water (e.g., Pluronic w ). In this method, the copolymer and the drug are directly dissolved in water 

or an aqueous solution such as phosphate-buffered saline (PBS) with subsequent heating and/or 

stirring to induce micellization. Successful loading of doxorubicin, etoposide, nystatin, and haloperidol 

into Pluronic w micelles has been achieved using this method. 33,34 

For copolymers and drugs with limited aqueous solubility, the evaporation method can be used 

to prepare the micelle formulations. This method involves first dissolving the copolymer and the 

drug in a common solvent or a mixture of two miscible solvents. The mixture is then stirred, and the 

solvent is allowed to evaporate, yielding a copolymer–drug film that can be reconstituted in warm 

water or a buffer to form the drug-loaded micelles. 35–38 

The most commonly used method for the preparation of drug-loaded micelles from poorly 

water-soluble copolymers is the dialysis method. The copolymer and the drug are first dissolved in a 

common organic solvent that is miscible with water. Slow addition of water results in the formation 

of micelles with swollen cores. The solution is then dialyzed against a large volume of water to 

remove the organic solvent. 3,28,32,39 As the percentage of solvent decreases, the degree of swelling 

in the micelle core decreases to form intact drug-loaded micelles. Alternatively, the solvent can be 

removed by evaporation of the solvent from the polymer–drug mixture to form micelles (i.e., 

emulsification method). 38,40,41 The selection of the organic solvent and the ratio of water-to-solvent 

employed have been shown to greatly affect the physical properties and drug loading of the 

micelles. 39 


324 

17.2.4 PROPERTIES OF MICELLE FORMULATIONS 

The drug loading and drug release profiles of block copolymer micelle formulations will influence 

the in vivo therapeutic efficacy of the system. There is a balance between the amount of drug that 

can be loaded into the hydrophobic core and the amount of drug that can be released from the same 

core. 36 The drug loading and drug release from the micelles is highly dependent on the partitioning 

behavior of the drug into the micelle. 

17.2.4.1 Drug Partitioning 

Drug loading into block copolymer micelles is achieved by the partitioning of the lipophilic agents 

into the micelle core. The partition coefficient (Kv) refers to the extent to which a drug selectively 

partitions into a micelle and can be expressed as follows: 

K v Z ½DrugŠ micelle 

½DrugŠ aqueous 

(17.1) 

where [Drug] micelle and [Drug] aqueous represent the molar concentrations of the drug in the micelle 

phase and aqueous phase, respectively. 23,42 Many different experimental methods including surface 

tension, 23 UV–vis spectroscopy, 43,44 fluorescence, 33,42,45,46 solubility in model solvents, 47,48 and 

theoretical methods 49,50 have been used to study the partitioning behavior of molecules in block 

copolymer micelle solutions. The length of the hydrophobic block influences both the CMC and the 

partition coefficient. An increase in either the length or the hydrophobicity of the core-block has 

been shown to improve the partitioning behavior of hydrophobic agents into the micelle core. 51,52 

Hydrophobic molecules such as pyrene preferentially reside in the hydrophobic micelle core. 50 

However, the micelle core is not the only site for drug solubilization. Various amphiphilic solubilizates 

can reside at the core–corona interface or in the corona itself. 45,47,53 

17.2.4.2 Factors Affecting Drug Loading 


The partitioning behavior of drugs into the micelle core, core–corona interface and/or corona 

controls the amount of drug that can be solubilized in the micelles. However, for a block copolymer 

micelle system to act as a true carrier, it must retain the drug within the micelles for prolonged 

periods of time following administration. Also, it is highly unlikely that a single block copolymer 

micelle system can effectively deliver all types of drugs. Therefore, selecting a core block for the 

drug of interest is an important consideration in the design of an effective micelle delivery system. 

Micelle drug loading is influenced by several factors, including the hydrophobicity and length of 

the core-forming block, the length of the corona-forming block, the specific interactions between 

the drug and core, and the method of micelle preparation. 

The hydrophobicity and length of the core-forming block have a great influence on the size of 

the micelles and the loading into the core of the micelle. For example, Lin et al. reported that 

increasing the degree of hydrophobicity of the core-forming block improved the drug loading into 

the micelles. 54 Also a larger amount of drug can be incorporated into micelles formed from 

copolymers having a larger core-forming block in comparison to a shorter hydrophobic block. 

For example, Shuai et al. demonstrated that for doxorubicin-loaded micelles formed from PEGb-PCL 

copolymers with a constant PEG block length, increasing the PCL block length resulted in 

an increase in doxorubicin loading from 3.1% to 4.3% (wt./wt.). The hydrophilic block length can 

also influence drug loading. By increasing the corona-forming block, the CMC will also increase, 

and this may result in a decrease in the aggregation number (N agg). For example, Gadelle et al. 



showed that increasing the PEG block length caused an increase in the CMC, a decrease in the Nagg, 

and a decrease in the degree of solubilization of hydrophobic agent into the micelles. 55 

Specific interactions between the core-forming block and entrapped drug can also affect drug 

loading in the micelles. This is referred to as polymer–drug compatibility where the compatibility is 

defined as interaction between the two species that does not involve any chemical change to either 

the drug or the polymer. Typically, the interactions that are present between polymer and drug are 

hydrogen bonding, ionic, and pi–pi. Lee et al. showed that by increasing the number of functionalized 

carboxylic acid (COOH) groups from 0% to 19.5% (weight percent of repeat units), they were 

able to augment papaverine loading from 3.5% to 15% (wt./wt.) in PEG-b-PDLLA micelles. They 

attributed the increase in papaverine loading levels to hydrogen bonding interactions between the 

COOH groups of the copolymer and the unpaired electron groups of the oxygen and/or nitrogen 

atoms in the drug. 56 Other approaches include improving drug loading by enhancing the local 

microenvironment of the micelle core through the means of conjugation of a small amount of 

drug 16 or conjugation of side chains with similar structures to the drug. 57,58 

Finally, the method that is selected for preparation of the micelles has also been shown to have 

an effect on the overall degree of drug loading. Yokoyama et al. have shown that using different 

micelle preparation techniques led to distinctly different drug loading efficiencies for 

campthothecin (CPT) in a series of PEG-b-poly(aspartate ester) copolymers. 38,59 The use of the 

dialysis method resulted in a CPT loading efficiency of 19% and a cloudy formulation with large 

aggregates and drug precipitates. In contrast, when CPT was loaded using the evaporation method, 

a loading efficiency of 73% was obtained and also a clear solution was observed. 38 In addition to the 

selection of the preparation method, method-specific parameters (nature of organic solvent, solvent 

ratio) have also been proven to affect drug loading into the micelles. Yokoyama et al. showed that 

the use of dimethylformamide to prepare KRN 5500 loaded PEG-b-poly(16-b-cetyl-L-aspartate-cob-benzyl-L-aspartate) 

micelles resulted in a higher drug loading level when compared to using 

dimethylsulfoxide. 60 The ability to achieve a high drug loading level is important as this allows for 

maximization of the drug to polymer ratio that, in turn, means that less copolymer will need to be 

administered per dose of formulation. 

17.2.4.3 Factors Influencing Drug Release from Polymeric Micelles 

A drug can be localized in potentially three different areas in a block copolymer micelle: in the core, 

at the core–corona interface, and at the corona. 53 The drug in the micelle core has a longer diffusion 

path relative to the drug that is located at the core–corona interface or the corona. The drug that is 

released from the core–corona interface or the corona is typically referred to as a burst release, an 

initial discharge of drug in a relatively short period of time (typically hours). For example, Bromberg 

et al. observed a burst release (5–11%) of b-estradiol from the poly(acrylic acid) (PAA) corona 

of Pluronic w -PAA micelles. 61 In most cases, after a burst release, a slower and more consistent 

release that typically lasts for a longer period of time (typically days to months) is observed. For 

example, Kim et al. demonstrated slow release of indomethacin (i.e., less than 30% released over a 

period of approximately 100 h) from Pluronic w /PCL micelles. 62 

The kinetics of drug release from micellar systems has been studied extensively by many 

groups. 36,45,61,63–65 Despite the extensive release studies, the mechanism of drug release from 

block copolymer micelles is not fully understood. However, many groups believe that diffusion 

and copolymer degradation are the two principal mechanisms of release provided that the micelles 

are stable under the release conditions investigated. 63 If the interaction between the polymer and the 

drug is strong and the rate of biodegradation is fast, then the degradation rate would govern the rate 

of drug release. However, the rate of drug release from micelles, as observed in vitro, usually 

exceeds the rate of copolymer degradation. As a result, polymer degradation can most often be 

ruled out as one of the main mechanisms of drug release. Therefore, diffusion of the drug may be 

considered the principal mechanism of release from micelles. 


326 


Drug release is influenced by three main factors: the characteristics of the drug (amount of drug, 

molecular weight of the drug, and molecular volume of the drug), the properties of the core-forming 

block (molecular weight of the core, nature of the core, and physical state of the core), and the 

degree of polymer–drug compatibility. 

The amount of drug present in the micelle has been found to influence the release. Most often it 

has been found that an increase in the concentration of drug present results in a decrease in the rate 

of drug release. 36,65–67 For example, Gref et al. reported that higher concentrations of lidocaine in 

PEG-b-poly(D,L-lactide-co-glycolide) (PEG-b-PLGA) micelles resulted in a slower release rate for 

the drug. 66 Similarly, Lim Soo et al. observed that increasing the amount of estradiol in PEG-b-PCL 

micelles decreased the rate of drug release. 65 

The physical properties of the drug, including molecular weight and molecular volume have 

also been shown to influence their rate of release from micelles. An increase in the molecular 

weight of the drug forces a greater reorientation of the polymer chains for movement of the drug 

molecules through the polymer matrix. An increase in the volume of the drug will also require a 

greater effort to reorganize the polymer to accommodate its movement. As a result, larger or 

heavier molecular weight drugs will tend to have lower diffusion coefficients as opposed to 

smaller molecular weight agents. 

The rate of diffusion of a drug in and out of the micelle is also greatly influenced by the 

properties of the micelle core. In general, increasing the molecular weight of the core-forming 

block will increase the size of the core provided that the hydrophilic block length is held constant. 

This increase in the molecular weight of the hydrophobic block and the size of the core has been 

found to decrease the overall rate of release rate of the entrapped agent. 63,68–71 

The nature of the core, specifically the hydrophobicity and/or polarity of the core, influences the 

release rate of the drug as it determines the permeability of the core to aqueous media. Micelles with 

a highly hydrophobic core will likely have less aqueous fluid in the core in comparison to micelles 

that have a more hydrophilic or polar core. Therefore, micelles prepared from more hydrophilic 

cores will likely have an accelerated rate of drug release. The physical state of the core will also 

affect the diffusion of drug from the micelles. Polymers that have a high T g or bulky groups present 

on their backbone are limited in terms of their ability to reorient. As a result, these polymers will 

form micelle cores with a high microscopic viscosity, resulting in a low diffusion coefficient for the 

drug that is incorporated because of limited movement. 72 This is in contrast to polymers that are 

more rubbery or gel-like that will tend to form cores that have a low microscopic viscosity and a 

higher rate of diffusion for the incorporated drugs. 

Finally, the degree of polymer–drug compatibility greatly influences the rate of drug release 

from the micelles. Drugs that are highly miscible with the core-forming block have a high degree of 

polymer–drug compatibility. It is expected that these drugs would be molecularly dissolved in the 

core and not exhibit any crystallinity such that the release would be relatively fast. However, it is 

necessary to consider that the actual effect of a high degree of miscibility on drug release depends 

on the strength and extent of the core–drug interactions. 36,45 An increase in the extent of interaction 

between a drug and the polymer will result in a decrease in the diffusion coefficient of the drug 

(slower release). 73 The ability to provide sustained release of the drug from micelle systems is 

particularly beneficial for their use in therapeutic applications. 

17.3 PHYSICAL CHARACTERISTICS OF POLYMERIC MICELLES FOR 

CLINICAL APPLICATIONS IN CANCER THERAPY 

In the past two decades, many groups have reported on the development and optimization of 

polymeric micelles for the delivery of anti-cancer drugs. In particular, much effort has focused 

on the formulation of doxorubicin and paclitaxel in micelles because of their dose-limiting toxicities 

in currently approved formulations and high potency against a wide variety of cancers. To 



(a) 

O 

OCH 3 O 

date, several doxorubicin- (i.e., NK911, SP1049C) and paclitaxel-loaded (i.e., Genexol-PM, 

NK105) micelle formulations have entered clinical trial development. 

17.3.1 DOXORUBICIN 

Doxorubicin (Adriamycin w ), a member of the anthracycline antibiotics family, is currently 

employed for treatment of many types of cancer because of its broad anti-tumor spectrum. It is 

approved, either alone or in combination with other chemotherapeutic drugs, for treatment of 

breast, ovarian, bladder, lung, thyroid, and gastric cancer as well as malignant lymphoma, acute 

forms of leukemia, and sarcomas. 74 The cytotoxic effect of doxorubicin is said to result from a 

combination of cell damage mechanisms, some that have not been conclusively identified. One of 

the effects of doxorubicin that is known to contribute to its anti-tumor activity is the inhibition of 

cellular replication and DNA transcription as a result of an irreversible change in the DNA tertiary 

structure. 75 This non-cell type specific cytotoxic effect of doxorubicin is mainly attributed to its 

ability to intercalate between DNA base pairs with its planar structure (Figure 17.3a). Indeed, the 

planar moiety of doxorubicin has been shown to improve the stability of micelles when loaded in 

the micelle core. 76 Doxorubicin has been successfully loaded into micelles formed from PEG-bpoly(3-caprolactone) 

(PEG-b-PCL), 41 PEG-b-poly(D,L-lactide-co-glycolide) (PEG-b-PLGA), 77,78 

PEG-b-poly(propylene oxide)-b-PEG (PEG-b-PPO-b-PEG; Pluronic w ), 26,79–81 and PEG-b-poly 

(aspartic acid) (PEG-b-PAsp). 16,32,82,83 Currently, two doxorubicin-loaded micelle formulations, 

NK911 (i.e., PEG-b-PAsp) and SP1049C (i.e., Pluronic w ) are being investigated in clinical trials. 

17.3.1.1 NK911 

H 3C 

OH 

OH 

O 

O 

OH 

O 

CH2OH OH 

NH 2 

NK911 is a doxorubicin-loaded PEG-b-PAsp copolymer micelle formulation developed by 

Kataoka and co-workers. Doxorubicin is both physically entrapped in the micelle core and chemically 

conjugated to the aspartic acid side chains of the core-forming block via amide linkages. 

NK911 is currently undergoing phase II clinical trial evaluation for efficacy and toxicity profiles. 

The formulation entered phase I in 2001, and the results from this trial revealed that NK911 exhibits 

a longer circulation half-life, a larger area under the curve (AUC), and reduced toxicities in 

comparison to the conventional formulation of doxorubicin. 82 The prolonged circulation lifetime 

of NK911 is attributed to its composition and the monomodal size distribution of the micelles with a 

mean diameter of approximately 50 nm. 16,82 

O 

O OH 

FIGURE 17.3 Structures of anti-cancer drugs, (a) doxorubicin and (b) paclitaxel, that have been formulated in 

polymeric micelle formulations for clinical trial development. 


(b) 

O 

NH 

OH 

O 

O 

O 

HO 

O 

O 

H 

O 

O 

O

328 

The amphiphilic or polar nature of doxorubicin (log P value of 0.52) facilitates the partitioning 

of this drug from the micelle core into the external medium. 84 Therefore, chemical conjugation of 

doxorubicin to the core-forming block can enhance drug retention and retard drug release from the 

micelles. The initial development of the PEG-b-Asp micelles for drug delivery consisted of only 

conjugated doxorubicin to the PAsp core. Specifically, PEG-b-PAsp copolymers were synthesized 

from PEG-b-PBLA with the subsequent addition of doxorubicin to the carboxyl groups on the side 

chains of the core-forming block. As a result of the conjugation, a 37 mol% substitution of doxorubicin 

was achieved that is equivalent to 20 mg/mL doxorubicin. 85 The doxorubicin residues of the 

conjugates stabilize the micelle core through pi–pi stacking and hydrophobic interactions, 

providing the micelles with excellent kinetic stability in vivo. 76 

Subsequent studies revealed that the chemically conjugated doxorubicin exerts negligible antitumor 

activity in vitro and in vivo as a result of the lack of a cleavable linkage between the drug and 

the copolymer. Therefore, the next generation of the PEG-b-PAsp micelles was prepared to include 

physically encapsulated doxorubicin within the micelle core as well as the chemically conjugated 

drug. The amount of doxorubicin that could be physically entrapped within the core was found to be 

proportional to the amount of conjugated doxorubicin present. The enhanced compatibility between 

the physically entrapped drug and the micelle core is attributed to the presence of the chemically 

conjugated drug. In vivo studies confirmed that micelles with the highest amount of physically 

entrapped doxorubicin had the most potent anti-cancer activity. 32 

The formation of a dimerized form of the physically entrapped doxorubicin was detected within 

the micelles. Despite the fact that the dimer was found to have insignificant anti-tumor activity 

in vitro and in vivo, it plays a supplementary role to stabilize the entrapped doxorubicin. The 

addition of dimers to the micelles as part of the physically entrapped components can further 

prolong the circulation lifetime of the micelles in vivo. 86 However, the addition of doxorubicin 

dimers was also found to decrease the aqueous solubility of the drug-loaded micelles that compromised 

the stability of the micelles following lyophilization and long-term storage. Therefore, the 

fully optimized NK911 formulation was prepared such that it only contained doxorubicin 

monomer. 82 

17.3.1.2 SP1049C 


SP1049C is a doxorubicin-loaded Pluronic w copolymer micelle formulation developed by Kabanov’s 

group and Supratek Inc. (Montreal, Canada). In this formulation, doxorubicin is physically 

encapsulated in micelles assembled from two different types of Pluronic w copolymers, L61 and 

F127 in a 1:8 w/w ratio. 26 In October 2005, SP1049C was granted orphan drug status by the FDA 

for the treatment of oesophageal carcinoma based on results from the phase I study; meanwhile, its 

phase II clinical trial results are currently under final review. In contrast to NK911, the phase I 

clinical trial of SP1049C deliberately recruited patients with cancers that were refractory following 

conventional treatment and patients with previous anthracycline treatment. 81 This is due to the 

established chemosensitizing effect of SP1049C against multi-drug resistant (MDR) cancers, an 

effect that is attributed to its Pluronic w copolymer composition. 

The Pluronic w copolymers are the class of copolymers that have been most widely evaluated 

and employed in pharmaceutical applications. Studies have indicated that Pluronic w copolymers 

having an hydrophile-lipophile balance (HLB) of less than 19 and a moderate molecular weight are 

most effective in overcoming MDR. 34 In a separate study, Pluronic w L61 that has an HLB value of 

7 was found to be the most effective chemosensitizer among other Pluronic w copolymers. 26,34 

Doxorubicin loaded in Pluronic w L61 micelles was found to be 2–3 orders of magnitude more 

toxic than conventional doxorubicin when evaluated in a range of drug-resistant cancer cells (e.g., 

MCF-7/ADR breast carcinoma, SKVLB ovarian carcinoma, KBV oral epithelial carcinoma, LoVo/ 

Dox colon adenocarcinoma, P388-Dox murine leukemia cells, SP2/0Dox myeloma as well as 

others). 26 Furthermore, mechanistic studies have shown that it is the unimers rather than the 



micelles that are responsible for the reversal of MDR. 26,34 More details on the mechanisms and 

effects of Pluronic w as chemosensitizers are given in Section 17.4. 

Despite the effectiveness of micelles formed from only Pluronic w L61, the hydrophobicity of 

the copolymer results in unfavorable stability issues. Phase separation and micelle aggregation 

leading to microembolic effects have been observed in preliminary in vivo toxicity studies of the 

L61 formulation. Kabanov and co-workers solved this problem by introducing a higher molecular 

weight and more hydrophilic Pluronic w copolymer (F127) into the L61 micelle formulation. By 

varying the ratio of the L61 to the F127 copolymer, it was found that at 0.25 w/w% L61 (i.e., 

effective dosing concentration of L61) and 2.0 w/w% F127, the Pluronic w micelles maintained their 

effective diameter of 22.4 nm with no aggregation. 26 In addition, the IC 50 of doxorubicin loaded in 

the L61/F127 micelles was similar to that of the L61 formulation. Therefore, the optimized 

SP1049C formulation utilized for the clinical studies contains 2.0 mg/mL doxorubicin, 2.5 mg/ 

mL L61, and 20 mg/mL F127 in 0.9% sodium chloride. 81 

17.3.2 PACLITAXEL 

Similar to doxorubicin, paclitaxel is a chemotherapeutic drug that has been loaded into block 

copolymer micelles for investigation in clinical trials. Paclitaxel is an anti-neoplastic drug that 

was first extracted from the bark of Pacific yew trees (Figure 17.3b). The cytotoxic effect of 

paclitaxel is mainly attributed to its ability to promote the assembly as well as prevent the depolymerization 

of microtubules. 87 Stabilization of the microtubule networks inhibits normal 

interphase and mitotic functions that, in turn, impedes cell division. Paclitaxel is approved for 

treatment of ovarian, breast, and non-small cell lung cancers as well as AIDS-related Kaposi’s 

sarcoma. 87 Paclitaxel is highly hydrophobic with a log P value of 3.96 and low aqueous solubility 

of 1 mg/mL. 88,89 The commercialized formulation of paclitaxel, known as Taxol w , contains the 

drug dissolved in a mixture of Cremophor EL and ethanol that increases its solubility by more than 

6000-fold. However, administration of Cremophor EL leads to severe dose-limiting toxicities such 

as hypersensitivity reactions and neuropathy. 90 As a result, much effort has gone into the development 

of a novel formulation of paclitaxel using biocompatible and biodegradable polymeric 

micelles. Paclitaxel has been successfully loaded into PEG-b-poly(D,L-lactide) (PEG-b- 

PDLLA), 35,37,90,91 poly(N-vinyl-pyrrolidone)-b-PDLLA (PVP-b-PDLLA), 28,92 PEG-b-PCL, 93–95 

and PEG-b-poly(d-valerolactone) (PEG-b-PVL). 95 Genexol-PM (i.e., PEG-b-PDLLA) and 

NK105 (i.e., PEG-b-PAsp) represent two paclitaxel-loaded polymeric micelle formulations that 

have entered clinical trial development. 

17.3.2.1 Genexol-PM 

Genexol-PM, developed by Samyang Co. (Korea), is a paclitaxel formulation based on PEG-b- 

PDLLA micelles. In 2004, results from a phase II clinical trial that evaluated Genexol-PM and 

cisplatin as treatment for patients with advanced gastric cancer was reported. 96 In the first clinical 

trial, the AUC and Cmax of Genexol-PM was found to be lower than that of Taxol w ; however, the 

maximum tolerated dose (MTD) was increased from 135 to 390 mg/m 2 . 90 Preclinical studies also 

revealed similar pharmacokinetic profiles for Genexol-PM and Taxol w in the B16 melanoma 

murine model. There are two explanations for the above observations. First, the liquid-like core 

of PDLLA acts as a solubilization site for paclitaxel as opposed to an actual drug carrier, 35 and 

second, the decrease in AUC and Cmax occurs as a result of the shift in the drug accumulation to the 

tumor tissues that was evidenced in preclinical biodistribution studies. 97 Formulation development 

of paclitaxel-loaded PEG-b-PDLLA micelles was mostly reported by Burt and co-workers. 35,37,91 

Genexol-PM is prepared by the evaporation method that involves co-dissolving paclitaxel and 

PEG-b-PDLLA in acetonitrile with subsequent evaporation of the solvent to form a copolymer– 

drug matrix. The matrix is then heated to 608C with the addition of water or buffer for self-assembly 


330 

of paclitaxel-loaded micelles. 97 The evaporation method is much more suitable than the direct 

dissolution method because paclitaxel and PEG-b-PDLLA are relatively insoluble in aqueous 

media. Indeed, the method employed for preparation of the paclitaxel-loaded PEG-b-PDLLA 

micelles can influence the extent of drug loading and physical stability of the formulation. The 

choice of solvent strongly influences the extent to which paclitaxel can be solubilized in PEG-b- 

PDLLA micelles as reported by Zhang et al. 37 Among many of the solvents tested, only acetonitrile 

resulted in a clear micelle solution with no evidence of drug precipitation. It was suggested that 

acetonitrile enhances the miscibility between paclitaxel and the PDLLA block that, in turn, reduces 

the possibility of phase separation during the solvent evaporation process. 37 

Burt and co-workers also studied the use of PEG-b-poly(D,L-lactide-co-caprolactone) and PEGb-poly(glycolide-co-caprolactone) 

copolymers for encapsulation of paclitaxel. The addition of 

hydrophobic caprolactone to PDLLA could potentially increase paclitaxel loading through the 

increased hydrophobicity of the core-forming block. However, these micelles were found to 

have poor physical stability and resulted in the precipitation of paclitaxel. 35 The PEG-b-PDLLA 

micelles can solubilize as much as 5% (w/v) paclitaxel and remain stable for up to 24 h. This is 

mainly attributed to the absence of crystalline paclitaxel in the paclitaxel/PEG-b-PDLLA matrix as 

revealed by x-ray diffraction studies. In fact, paclitaxel is molecularly dissolved in the PDLLA core 

that prevents drug precipitation and results in excellent solubilization of the drug. 37,91 

The amorphous state of the paclitaxel–PDLLA core can be problematic for in vivo applications. 

Studies have indicated that the presence of crystalline drug in the core can retard drug release. 69,98 

Therefore, the amorphous core of the paclitaxel-loaded PEG-b-PDLLA micelles can lead to a much 

faster drug release profile. As demonstrated by Burt et al. following intravenous administration 

paclitaxel rapidly dissociates from the micellar components in the circulation. 35 Therefore, it is 

likely that Genexol-PM acts as a solubilizer for paclitaxel rather than as a true drug carrier. 

17.3.2.2 NK105 


In 2005, Kataoka and co-workers, in collaboration with NanoCarrier Co., Ltd. and Nippon Kayaku 

Co., Ltd., first reported on NK105, a paclitaxel-incorporated micelle formulation. 99 The phase I 

clinical trial of NK105 first began in April 2004. Similar to NK911, the building block of NK105 is 

PEG-b-PAsp copolymers but with a slight modification to the aspartate blocks such that the 

carboxyl residues are converted into 4-phenyl-1-butanolate by an esterification reaction. The 

modification increases the hydrophobicity of the core-forming block that results in enhanced 

entrapment and retention of paclitaxel in the micelles. The resulting drug-loaded micelles 

contain 23% (w/w) paclitaxel with a weight–average diameter of approximately 85 nm. 

With the improved micelle core stability, NK105 is predicted to act as a true drug carrier rather 

than a solubilizer. As indicated by the in vivo preclinical studies, NK105 is extremely stable in vivo 

with a 90-fold increase in the AUC and a 4–6-fold increase in the terminal phase half-life. 99 In 

addition, the tumor AUC was found to be 25-times higher than that of Taxol w . The superior in vivo 

anti-tumor activity of NK105 in a human colorectal cancer xenograft model (HT-29) was a result of 

the enhanced drug accumulation in the tumor, allowing for more drug-loaded micelle–cell 

interactions. These results highlight the importance of polymer–drug compatibility and core stability 

in the design of effective micellar delivery systems. Therefore, the recent studies that focus on 

customizing copolymer compositions to accommodate the distinct properties of different drugs 

such as camptothecin, 38,100 KRN 5500, 60 and cisplatin 101,102 may lead to the development of true 

polymeric micelle carriers for delivery to tumor tissues. 

17.4 INTERACTIONS OF POLYMERIC MICELLES WITH CANCER CELLS 

Once the drug-loaded micelles have reached the tumor site, there are several mechanisms that 

have been proposed for the delivery of micelle-incorporated drugs into the cancer cells as 



(a) 

(b) 

illustrated in Figure 17.4. 103–106 In general, drugs can enter the cells via diffusion as free 

molecules following release from the micelles (Figure 17.4a and b) or be internalized as drug 

molecules entrapped within the micelles (Figure 17.4c–f). In addition, certain block copolymers 

have also been found to act as chemosensitizers or modulators of drug efflux pumps. 26,34,107–111 

17.4.1 DRUG DIFFUSION 

Lysosome 

Nucleus 

Lysosome 

Endosome 

FIGURE 17.4 Proposed mechanisms for the delivery of micelle-formulated drugs into cancer cells. (a) Drug is 

released from the micelles and diffused into the cell as free drug. (b) The drug-loaded micelles diassemble in 

the extracellular space and release the drug; the released drug then diffuses across the cell membrane into the 

intracellular space. (c) Intact drug-loaded micelles are internalized via adsorptive pinocytosis after adsorption 

onto the cell surface or (d) by fluid-phase endocytosis. The internalized drug-loaded micelles eventually reach 

the lysosome where the drug can be released (e) by diffusion or (f) following disassembly or degradation of the 

micelles. 

A free drug can diffuse across the cell membrane following disassembly of the micelles or release 

of the drug from intact micelles. The extent and rate of diffusion of the free molecules across the 

cell membrane will depend largely on the physico-chemical properties of the drug (water solubility, 

charge, molecular weight, and partition coefficient). Small hydrophobic molecules can easily 

permeate the membrane and enter the intracellular space. Studies have shown that the rate of 

cell uptake of a free drug (non-micelle-incorporated) is faster than that of drug incorporated in 

micelles. 103–105 In this way, the release profile of the drug from the micelles can have a significant 

influence on the rate of cellular uptake of the drug. As the drug is released from the micelles, there is 

an increase in the gradient between the concentrations of free drug on the exterior and interior of the 

cell that, in turn, drives diffusion of the drug into the intracellular compartment. A micelle system 

with a slow drug release profile will only have a limited pool of free drug available for diffusion into 

the cell. For example, Maysinger et al. demonstrated that for PEG-b-PCL micelles loaded with 

benzo[a]pyrene that has an apparent partition coefficient (Kv) of 690, the rate of cell uptake of the 

micelle formulated probe was similar to that of the rate of uptake of free benzo[a]pyrene. In 

contrast, for micelles loaded with DiI that has a Kv of 5800, a much lower degree of cell uptake 

of the micelle formulated probe was observed when compared to free DiI. 45,103 Therefore, at the 

site of interest, an increase in the rate of drug release from the micelles can result in a greater extent 


(e) 

(c) 

(f) 

(d)

332 

of intracellular drug accumulation if the drug can easily diffuse across the cell membrane. Micelle 

systems that respond to external stimuli as a means to promote drug release at the tumor have also 

been designed and are discussed further in Section 17.6. Another route for the entry of micelleformulated 

drug into cells includes the internalization or endocytosis of the drug-loaded 

micelles. 103,104,106 

17.4.2 CELLULAR INTERNALIZATION 

Macromolecules and nanosized particles including polymer–drug conjugates, liposomes, and 

micelles are unable to diffuse through the cell membrane into the intracellular compartment. In 

order to enter the cell, these drug delivery vehicles require an internalization mechanism such as 

phagocytosis or pinocytosis. Only certain cell types called phagocytic cells (macrophages, neutrophils) 

can undergo phagocytosis. Particulate materials with diameters of less than 1 mm can be 

internalized into these cells in membrane-bound vesicles called phagosomes. 112 Conversely, pinocytosis 

is known to be non-cell type specific. This mechanism occurs by invagination of the cell 

membrane that is similar to the mechanism in phagocytosis. However, the pinocytic vesicles that 

pinch off from the cell membrane are generally smaller than phagosomes and have diameters 

ranging from 100 to 200 nm. 112 The pinocytosis mechanism can be further categorized into 

fluid-phase and adsorptive. 

Fluid-phase pinocytosis involves the engulfment of soluble materials and extracellular fluid 

during continual cell membrane ruffling. This is in contrast to adsorptive pinocytosis that involves 

the association of particulate materials with cell membrane components prior to internalization. 113 

(e) 

(a) 

k d 

k a 

(d) 

Early endosome 

(b) 

(f) 

Late endosome 


(c) 

Nucleus 

FIGURE 17.5 Internalization of ligand-conjugated micelles via receptor-mediated endocytosis. (a) The 

ligand-coupled micelles approach the cancer cells and (b) bind to the cell surface receptors depending on 

the association constant (ka) and the dissociation constant (kd) of the receptor–ligand complexes. (c) The 

micelles are then rapidly internalized into the cells via receptor-mediated endocytosis. (d) The internalized 

receptor and the ligand-conjugated micelles then undergo endosomal sorting in the early endosomes. In most 

cases, (e) the receptor is recycled to the cell surface, and (f) the micelles enter the late endosome that eventually 

fuses with the lysosome. 



For example, micelles that are conjugated with specific ligands can bind to the corresponding 

receptors that are expressed on the cell surface; as a result, the micelles can be internalized via 

receptor-mediated endocytosis. 78,114,115 More details on receptor-mediated endocytosis are 

provided in Section 17.6 and Figure 17.5. In most cases, micelles are internalized into the cells 

through fluid-phase endocytosis. 77,116,117 

Fluid-phase endocytosis is expected to result in a lower degree of cell uptake in comparison to 

adsorptive pinocytosis or diffusion of free hydrophobic agents. Differences in the cell uptake of 

micelle-incorporated agents and non-micelle-incorporated agents have been demonstrated in 

various studies. For example, using a hydrophobic fluorescent probe as a tracker for micelles, 

the rate and amount of intracellular fluorescence was found to be much lower than incubation 

with free probes. 104 Thishasalsobeenconfirmedbyotherstudies that relied on confocal 

microscopy as a qualitative tool. 103,105 Rapoport et al. also demonstrated that the incorporation 

of doxorubicin and ruboxyl into Pluronic w micelles resulted in a decrease in the extent of cell 

uptake of the agents in ovarian cancer cells. 118 In a drug accumulation study at copolymer concentrations 

above the CMC of the Pluronic w copolymers, the accumulation of the fluorescent probe 

R123 was substantially decreased in both non-multi- and multi-drug resistant cancer cells. It was 

suggested that the R123-loaded micelles were transported into the cells via a vesicular route and 

followed accumulation kinetics different from that of the non-micelle incorporated R123. 34 

Endocytosis is known to be an energy-, time-, and pH-dependent process. 113 Many pharmacological 

manipulations have been employed in order to demonstrate that micelles are internalized 

via an endocytotic pathway. For example, a low pH and incubation at 48C resulted in the reduction 

of the internalization of micelles. 104 Also, pre-treatment of cells with metabolic inhibitors such as 

sodium azide and 2-deoxyglucose (DOG) that deplete the energy source for membrane ruffling 

resulted in a decrease in the extent of uptake of radiolabeled-drug loaded-PEG-b-PCL micelles by 

at least 80%. Another study conducted by Luo et al. reported similar findings, yet in this case, the 

micelles were tracked using fluorescently labeled-copolymer. 106 Furthermore, the pre-treatment of 

cells with Brefeldin A, known to alter the morphology of endosomes, was also found to reduce the 

extent of uptake of micelles. This suggests that an endosomal compartment may be involved in the 

internalization of the micelles that further supports the claim that micelle internalization does 

proceed by an endocytotic mechanism. 104 

In order to confirm that micelles have been internalized via the endocytotic pathway, different 

methods have been developed for the detection of micelles in the relevant organelles. For example, 

a pH-sensitive fluorescent probe has been used to confirm the presence of the micelles in acidic 

organelles using flow cytometry. 118,119 The subcellular localization of micelles can also be visualized 

using confocal microscopy by fluorescently labeling the micelles and counterstaining the 

intracellular organelles. 41,105,115 Hydrophobic fluorescent probes physically entrapped in micelles 

can also be used to track the micelles. The use of these probes to follow the pathway or fate of the 

micelles requires virtually no or minimal release of the probe during the time frame of the experiments. 

However, cell culture media often contain protein that can act to accelerate drug release. 120 

Therefore, the accuracy associated with using physically entrapped probes to track micelles is 

questionable because of the different subcellular localization patterns of free and micelle-incorporated 

drugs. 41,105 Ideally, fluorescent probes that are chemically conjugated to the copolymers 

provide the most reliable tool for following the micelles in vitro. Although copolymers also 

exist in unimeric form, the concentration of unimers in a micelle solution does not exceed the 

CMC of the copolymer and represents only a minor population. 106 Therefore, it is reasonable to 

assume that the fluorescence detected corresponds to the micelle population provided that the 

conjugation of the probe to the copolymers is stable. 

Internalized micelles have been shown to localize mainly in cytoplasmic compartments, in 

particular, the accumulation of micelles in acidic organelles (endosomes or lysosomes) has been 

observed. 105,118,119 Rapoport et al. have reported on the intracellular uptake and trafficking of 

Pluronic w micelles. Their results suggested that after the micelles were internalized in acidic 


334 

organelles, the non-ionic Pluronic w copolymers induced permeabilization of the membranes, 

resulting in the release of the encapsulated drugs into the cytosol or into the nucleus via diffusion 

of the drug. 118 The nuclear pore complexes on the nucleus allow for passive diffusion of small 

molecules that have diameters of less than 9 nm, whereas macromolecules as large as 39 nm can be 

actively transported into the nucleus. 121 Theoretically, micelles that are greater than 9 nm cannot be 

translocated into the nucleus unless nuclear targeting signals are present on the particles. For 

example, in cancer cells that have been incubated with tetramethylrhodamine-5-carbonyl azide 

(TMRCA)-conjugated PEG-b-PCL micelles for 24 h, only cytoplasmic localization of the micelles 

was observed, and the nuclear compartment was devoid of micelles. 105 Similar results were 

obtained by Shuai et al. with doxorubicin-loaded PEG-b-PCL micelles that have a delayed 

release profile for doxorubicin from the micelles. The authors demonstrated that following incubation 

of cells with micelle formulated doxorubicin for 2 and 24 h, the drug was only localized in 

cytoplasmic compartments, whereas incubation of free drug with cells resulted in nuclear accumulation. 

41 To date, the cellular internalization pathway and fate of micelles and micelle-encapsulated 

drugs as a function of time have not been clearly elucidated. However, the in vitro pathway and fate 

of micelles can have a significant impact on the cytotoxic effects of the formulated drug. The 

various anti-cancer agents that have been traditionally formulated in micelles have different sites of 

action within the cell. For example, paclitaxel and doxorubicin are required to bind to microtubules 

in the cytoplasm and DNA in the nucleus, respectively, in order to exert their cytotoxic effects. 

Therefore, if micelles could be engineered or designed to achieve a specific subcellular localization 

that matched the required localization of the drug, the cytotoxicity of the drug could be enhanced. 

17.4.3 IN VITRO CYTOTOXICITY 


The subcellular localization and the release profile of drug-loaded micelles can impact the cytotoxic 

effects of the drugs. In most cases, free drugs have been found to have lower IC 50 values than drugs 

formulated in micelles. Such an increase in the IC 50 of the micelle-formulated drugs is likely 

because of the delayed release of drugs from the micelle core. 41,95,122 

Shuai et al. reported delayed cell death in MCF-7 breast cancer cells when using a doxorubicin 

formulated PEG-b-PCL micelle system in comparison to free doxorubicin. 41 The confocal 

microscopy images showed that the micelle-incorporated doxorubicin resides in the cytoplasm, 

and free doxorubicin accumulates quickly in the cell nucleus. The delayed cell death could be due 

to the difference in the subcellular localization of free and micellized drug because doxorubicin can 

only exert its cytotoxic effect after reaching the nucleus. 41 Interestingly, as reported by Yoo and 

Park, doxorubicin-conjugated PEG-b-PLGA micelles, led to a 10-fold increase in the cytotoxicity 

of this drug. However, in this case, doxorubicin was found to localize in both the cytoplasm and the 

nucleus when delivered by micelles. 77 The discrepancy between the intracellular localization of 

doxorubicin formulated in the two distinct systems may be explained by their different drug release 

profiles. The PCL core is more hydrophobic than the PLGA core and can, therefore, provide better 

drug retention. Although doxorubicin was conjugated to the PLGA system that should result in a 

much slower release, the hydrolytic linkage is susceptible to cleavage and results in a faster release 

of the drug when compared to the PCL system. The cytotoxicity observed in the PEG-b-PLGA 

micelle system can be attributed to the doxorubicin that has been cleaved from the copolymer and 

released from the micelles because the conjugated drug was found to be non-cytotoxic. 16,32 

Furthermore, using a thermosensitive micelle system, Liu et al. obtained accumulation of 

doxorubicin in the cytoplasm and the nucleus following a temperature increase, and no nuclear 

accumulation of the drug was observed prior to the increase in temperature. 122 The delayed cytotoxicity 

of micelle-formulated drugs may seem to be a limitation in terms of the utilization of this 

technology as cancer therapy. However, the retention of drug in the micelles in vivo is required in 

order to ensure successful delivery of the drug to the tumor site. The relationship between drug 

retention in micelles and in vivo efficacy is discussed in Section 17.5. 



Copolymer micelles have also been found to amplify the effect of chemotherapeutic drugs in 

cancer cells. In the case of multi-drug-resistant cancer cells, certain copolymers can also act as 

chemosensitizers and increase the intracellular accumulation of the drug. The Pluronic w copolymers 

are the family of materials that have been investigated most extensively in terms of their role 

as chemosensitizers or modulators of drug efflux. 

17.4.4 INFLUENCE OR EFFECT ON MULTI-DRUG RESISTANCE 

Multi-drug resistance (MDR) is one of the most significant obstacles in chemotherapy and is often 

found in refractory cancers. MDR may be a result of the undesired efflux of a drug from the 

intracellular compartment. Transporter proteins that are overexpressed on the cell membrane of 

certain cancer cells are responsible for drug efflux. The drug efflux protein that has been studied 

most extensively is the ATP-dependent P-glycoprotein (P-gp). 26,123 This drug transporter is also 

expressed in healthy tissues such as the epithelial cells of the intestine and the endothelial cells in 

the blood–brain barrier. The normal physiological function of P-gp is to regulate molecular transport 

and provide protection or detoxification of the cell. 124 The normal physiological levels of drug 

efflux protein also function to limit the bioavailability of drugs delivered via the oral route as well as 

to limit drug delivery to the brain. 

Kabanov’s group has made significant contributions in the development of Pluronic w micelles 

as a tool for enhancing drug delivery both across the blood–brain barrier and into MDR 

tumors. 7,26,34,108,110,124–127 Their work in this area originated from the finding that doxorubicin 

formulated in Pluronic w micelles was more cytotoxic than free doxorubicin in drug-resistant cells. 

Pluronic w can increase the cytotoxicity of anthracyclines in drug-resistant cancer cells by 2–3 

orders of magnitude. For instance, the IC50 of daunorubicin in the presence of Pluronic w copolymers 

was decreased by 700-fold in resistant ovarian cancer cells (SKVLB). 125 Additional results on 

the cytotoxicity of anti-cancer drugs formulated in Pluronic w micelles in resistant cancer cells are 

summarized elsewhere. 108,125 Further mechanistic studies have demonstrated that the ability of 

Pluronic w copolymers to inhibit drug efflux proteins depends on their composition. 

Specific compositions of Pluronic w copolymers were evaluated in a screening study in order to 

demonstrate that the inhibitory effects of the copolymers on drug efflux were dependent on the 

length of the PEG and PPO segments. It was found that hydrophobic Pluronic w copolymers with 

values of HLB!19 and with an intermediate length of the PPO segments such as L61 (PEG2– 

PPO30–PEG2) and P85 (PEG26–PPO40–PEG26) are the most effective at inhibiting P-gp. 34 

SP1049C is a Pluronic w -based micelle formulation that has been selected for further preclinical 

development and has entered phase II clinical trial for the treatment of relapsed cancers. L61 was 

chosen as the base component for the SP1049C formulation with the addition of F172 (another high 

molecular weight hydrophilic Pluronic w ) copolymers to stabilize the formulation as discussed in 

Section 17.3. 26 

In many of the studies, the P85 copolymer was fundamental in the investigation of the 

mechanisms of Pluronic w copolymers for overcoming drug resistance in cancer. 23,108,109,124 To 

date, several mechanisms have been hypothesized to account for the role that Pluronic w copolymers 

play as drug efflux modulators, and these have been reviewed in detail elsewhere. 26,123,125 It has 

been confirmed that the unimers, as opposed to the copolymer micelles, are responsible for inhibiting 

drug efflux. For example, accumulation and cytotoxicity of doxorubicin were demonstrated 

to increase with an increase in the Pluronic w concentration up to the CMC of the copolymer in drugresistant 

human breast cancer cells (MCF-7/ADR) as well as head and neck cancer (KBv) cells. 

However, after reaching the CMC of the copolymers (the concentration when micelles are formed), 

both the intracellular doxorubicin content and the cytotoxicity were found to level off and/or 

decrease. 34 Similarly, the extent of intracellular accumulation of fluorescein, a multi-drug resistance-associated 

protein (MRP) substrate, in an MRP-expressing human pancreatic adenocarcinoma 

cell line (Panc-1) was found to be dependent on the concentration of Pluronic w present. 127 


336 


The ability of Pluronic w unimers to inhibit drug efflux transporters is mainly attributed to ATPdepletion. 

26,109 Use of certain metabolic inhibitors such as sodium azide can inhibit P-gp and result 

in increased drug accumulation by depleting the energy source for efflux across the cell membrane. 

Similar effects are obtained with the addition of Pluronic w copolymers to the cells as they result in 

an abolishment of the energy supply within the MDR cells. 34,109,125,127 For example, P85 and P105 

were found to reduce the activity of the electron transport chain reactions in the mitochondria in 

HL-60 cells, decreasing the degree of oxidative metabolism, indicating that Pluronic w copolymers 

can interrupt the production of ATP. 128 In addition, cell exposure to P85 has been found to induce 

reversible transient ATP depletion in Jurkat lymphoma T-cells. 129 These effects are much more 

profound in MDR cells because the drug efflux mechanisms are ATP-driven. 109 It was believed that 

MDR cells consume energy at a much higher rate than non-MDR cells; therefore, their activity is 

more susceptible to inhibition via ATP-depletion. As a result, the Pluronic w copolymers were 

capable of chemosensitizing the MDR cells to a greater extent than non-MDR cells. All of the 

above suggest that Pluronic w -induced ATP depletion is the major mechanism that is responsible for 

inhibiting drug efflux in MDR cells. 

Another proposed mechanism of drug efflux inhibition is membrane fluidization caused by 

Pluronic w copolymers. Non-ionic surfactants have been shown to increase membrane fluidity and 

contribute to the inhibition of P-gp efflux, resulting in the enhanced cytotoxicity of anthracyclines 

in MDR cells. 125,130 However, Pluronic w cannot increase the intracellular accumulation of P-gp 

substrate in MRP-expressing cells, indicating that membrane fluidization on its own is insufficient 

to inhibit drug efflux in drug-resistant cells. It was believed that the membrane fluidization effect 

works synergistically with the ATP-depletion effect induced by Pluronic w copolymers to chemosensitize 

MDR cells. 127 

Recently, Burt’s group reported on the synthesis of a series of low molecular weight PEG-b- 

PCL for inhibition of P-gp efflux in human colon adenocarcinoma (Caco-2). Interestingly, at 

concentrations above the CMC, the PEG-b-PCL copolymers were found to increase the accumulation 

of P-gp substrate, and little activity was observed at concentrations below the CMC. 111 

Nevertheless, the inhibition of P-gp efflux was attributed to the unimers, and the inhibition 

mechanisms of the PEG-b-PCL copolymers were said to be similar to that of the Pluronic w 

copolymers. 107 These low molecular weight PEG-b-PCL copolymers are a class of materials 

with potential applications in the delivery of drugs to MDR cancer cells or the formulation of 

drugs requiring improved bioavailability. 

17.5 IN VIVO BIOLOGICAL PERFORMANCE OF BLOCK COPOLYMER MICELLES 

In order to successfully reach and interact with the tumor cells, the micelle-formulated drug must 

first circumvent the MPS. As systemic drug delivery vehicles, block copolymer micelles may serve 

as either solubilizers and/or true drug carriers depending on their capability for drug retention and 

stability in vivo. Several of the micelle formulations that have been developed for hydrophobic 

drugs have been shown to provide a significant increase in their water solubility. 36 The use 

of biocompatible and biodegradable copolymers to form these micelles may result in formulations 

with reduced toxicity in comparison to formulations formed from some conventional surfactants 

or cosolvents. As a result, these copolymers may allow for higher doses of the drug to be administered, 

and they result in improved efficacy when compared to a conventional formulation of 

the drug. 

Micelles that function merely as solubilizers increase the aqueous solubility of the drug, but 

they do not retain the drug for prolonged periods of time following systemic administration. In this 

way, the pharmacokinetics and biodistribution profile of the drug may remain largely unchanged. 

By contrast, micelles may function as true carriers that are able to retain the drug until reaching the 



Endothelial cells 

Red blood cells 

Proteins 

Macrophages 

target or desired site. 16,99,101 Micelles that function as carriers can result in significant improvements 

in the toxicity profile and/or therapeutic efficacy of the encapsulated agents. 

The ability of a micelle system to function as a solubilizer or a true carrier depends on the 

following two factors: the stability of the micelles in vivo and the drug release profile in the 

presence of various blood components. The stability of block copolymer micelles in vivo 

depends on both thermodynamic and kinetic components as discussed in Section 17.2. 

17.5.1 FATE OF POLYMERIC MICELLES IN VIVO 

(a) 

(b) 

(c) 

Tumor interstitium 

MPS clearance 

FIGURE 17.6 In vivo fate of drug-loaded micelles following administration into the bloodstream. (a) Proteins 

can be adsorbed onto the micelle surface, leading to eventual clearance by the mononuclear phagocytic system 

(MPS). (b) The drug-loaded micelles can also circulate in the bloodstream, accumulate in the tumor interstitium 

via the enhanced permeability and retention (EPR) effect, and subsequently release the encapsulated 

drug to the proximal tumor cells. (c) In the case where the micelle-encapsulated drug has a high affinity for 

blood components, the drugs may be released from the micelle core at an accelerated rate to bind to proteins 

that are present in the circulation prior to reaching the tumor. 

The MPS serves as one of the major mechanisms for the elimination of nanosized drug delivery 

systems such as liposomes, nanoparticles, and micelles. 131–133 Clearance by the MPS is first 

mediated by an array of blood components that interact with the delivery vehicle or other 


338 

foreign bodies present within the systemic circulation, a process termed opsonization. 134–136 

Following intravenous injection, the micelles quickly distribute into various tissues and organs 

by way of the blood and lymphatic system. The primary elimination route for micelles is postulated 

to include uptake by the liver and spleen followed by clearance through the kidneys via the 

excretion of urine 35,76,137–139 and/or hepatobiliary excretion in the feces. 131,140 In this process, 

various plasma proteins play an important role because they adsorb to the surface of a delivery 

vehicle within the first few minutes of exposure, especially if the surface is charged or hydrophobic 

(Figure 17.6a). 141 The protein adsorption could cause opsonization, leading to rapid clearance of 

both carrier and encapsulated drug by the MPS. 10,142,143 As a result, a reduction in the extent of 

protein adsorption to delivery vehicles such as micelles has been considered to be one of the key 

strategies to increase the circulation lifetime of the vehicles in vivo. 

Various efforts to prolong the lifetime of micelles in vivo have been explored, including 

optimization of micelle size, size distribution, and surface properties. To this point, the micelles 

explored for applications in drug delivery have mostly been formed from copolymers that include 

PEG as the hydrophilic block. The presence of PEG at the surface of a delivery vehicle is known to 

impart stealth properties. 1,2,10 Good correlations have been found between the length of the PEG 

block and the extent of stabilization provided. 137 The degree of stabilization is largely determined 

by the thickness of the PEG layer. The PEG chains are present in a coil-like conformation on the 

surface of the micelles. The thickness of the surface layer (dh) may be calculated from the PEG 

block length using the following equation: dhZa 0.84 where a is the number of PEG units. 144 In a 

study by Gref et al. in Balb/C mice, a significant increase in the circulation lifetime of PEG-b- 

PLGA nanospheres was observed when the length of the PEG block was increased from 0 to 

20,000 g/mol, and the length of the PLGA block was held constant. The increased thickness of 

the protective PEG layer likely provides complete coverage of the more hydrophobic PLGA 

components. In the same study, the increase in the PEG length also resulted in a significant 

reduction in the extent of liver uptake of the particles that provided further evidence that PEG 

could effectively decrease the extent of opsonization and, therefore, limit MPS elimination. 66 

For materials to be eliminated by glomerular filtration through the kidneys, they must have a 

molecular weight of less than 50,000 g/mol. 17 The total molecular weight of a block copolymer 

micelle is typically on the order O200,000 g/mol, and the molecular weight of the individual 

copolymers ranges between 3000 and 25,000 g/mol. 145 Therefore, it may be postulated that only 

the copolymer single chains rather than intact micelles are eliminated by renal excretion. 

17.5.2 IN VIVO DRUG RELEASE PROFILES 


Micelles introduced into the circulation are exposed to a vast array of proteins, blood cells, and 

other blood components. A high affinity between the drug and major blood components commonly 

results in an acceleration of drug release from the micelles in vivo that, in turn, leads to a rapid 

elimination of the drug from the body (Figure 17.6c). 16 Studies have shown that the in vivo release 

profile of drug from micelles is largely determined by the compatibility between the drug and the 

core-forming block as well as the affinity between the drug and plasma proteins. 36,120,146,147 

Ramaswamy et al. investigated the distribution of paclitaxel in vitro in human plasma samples 

following incubation of free drug or drug incorporated in PEG-b-PDLLA micelles with the plasma. 

In this study, similar plasma distribution profiles were obtained following incubation of free and 

micelle-formulated paclitaxel with lipoprotein deficient and lipoprotein rich fractions. 147 In this 

way, this study clearly illustrates the extent to which the in vivo retention of drugs in colloidal 

carriers may be determined by the drugs’ affinity for plasma components. Therefore, the microenvironment 

within the delivery system must have the capability to compete with the various 

plasma components in order to achieve prolonged drug retention within the micelles in vivo. 



17.5.3 TOXICITY, PHARMACOKINETICS, AND ANTI-CANCER EFFICACY OF DRUGS FORMULATED 

IN BLOCK COPOLYMER MICELLES 

As previously discussed, various micelle-based formulations have been demonstrated, in both 

preclinical and clinical studies, to decrease the toxicity and increase the efficacy of the encapsulated 

agents. 16,82,90,148 To date, two of the drugs that have been most commonly formulated in block 

copolymer micelles are doxorubicin and paclitaxel (Figure 17.3). A general description and overview 

of both the doxorubicin and paclitaxel formulations in clinical development has been given in 

Section 17.3. At present, there are two micelle formulations of doxorubicin (i.e., SP1049C and 

NK911) and two of paclitaxel (Genexol-PM and NK105) that are in clinical trial development. 

17.5.3.1 Doxorubicin Formulations (SP1049C and NK911) 

SP1049C includes doxorubicin formulated in L61 and F127 Pluronic w copolymers at a ratio of 1:8 

(wt/wt). 26 For doxorubicin administered in the SP1049C formulation at a dose of 10 mg/kg, the 

AUC in healthy mice was found to be 14.6 mg h/mL while the maximum concentration (Cmax) in 

plasma was 4.47 mg/g. For doxorubicin administered as free drug, the AUC is 7.1 mg h/mL and the 

Cmax in plasma is 2.77 mg/g. In this way, the SP1049C formulation provides a 2.1-fold increase in 

the AUC. 

NK911 is another micelle formulation of doxorubicin and includes both chemically and physically 

entrapped doxorubicin within PEG-b-PAsp micelles. In its phase I trial, the administration of 

the NK911 formulation at a dose of 50 mg/m 2 was shown to provide a significant increase in the 

half-life (t1/2) and the AUC as well as a significant decrease in the steady state volume of distribution 

(Vss). Specifically, for doxorubicin administered to humans in NK911, the t1/2 values were 

reported to be t1/2,aZ7.5 min, t1/2,bZ2.8 h, and t1/2,gZ64.2 h, and the AUC was found to be 

3262.7 ng h/mL, and the Vss was 14.9 L/kg. In comparison to free doxorubicin, the t1/2 values 

were reported to be t 1/2,aZ2.4 min, t 1/2,bZ0.8 h, and t 1/2,gZ25.8 h, and the AUC was found to 

be 1620.3 ng h/mL, and the V ss was 24 L/kg. 82 In this way, the encapsulation of doxorubicin in the 

NK911 formulation increases the AUC and plasma Cmax for this drug by 36.4 and 28.6-fold, 

respectively. The change in the pharmacokinetic profile of the drug translates into changes in the 

efficacy and toxicity of the encapsulated agent. Specifically, the anti-tumor activity was evaluated 

in four tumor xenograft models in mice, namely, Colon 26, M5076, P388, and Lu-24. For the 

groups of animals receiving treatment with the micelle-formulated drug, a significant percentage of 

mice bearing colon 26 and MX-1 tumors were cured, whereas no cures were found for mice who 

received the same dose of free doxorubicin. Moreover, in all tumor models, treatment with the free 

drug resulted in more severe toxicity than that observed in animals receiving even the highest dose 

of NK911. 

17.5.3.2 Paclitaxel Formulations (Genexol-PM and NK105) 

Similar to doxorubicin, there are two promising formulations of paclitaxel that are under clinical 

trial evaluation: Genexol-PM and NK105. The main pharmacokinetic parameters of paclitaxel 

administered as Genexol-PM at a dose of 135 mg/m 2 are 5473 ng h/mL for the AUC and 12.7 h 

for the t 1/2b; whereas for paclitaxel administered as Taxol w , the AUC is 9307.5 ng h/mL, and the 

t1/2Z20.1 h. 90,149 In this way, the Genexol-PM formulation does not improve the pharmacokinetic 

profile of paclitaxel in vivo; therefore, the micelles in this formulation function primarily as 

solubilizers. In fact, it has been found that the drug is released quite rapidly from the PEG-b- 

PDLLA micelles following administration. However, as a result of the low toxicity of the micelles 

in comparison to Cremophor EL (the primary excipient in Taxol w ), the advantage of this formulation 

is the significant improvement in the maximum tolerated dose (i.e., MTDZ390 mg/m 2 ) 

when compared to the MTD of paclitaxel administered as Taxol w (i.e., 135–200 mg/m 2 ). 90,149 

In this case, the increase in the MTD is attributed to the biocompatibility or non-toxicity of the 


340 

PEG-b-PDLLA copolymers. Therefore, an improvement in the anti-cancer efficacy may be 

expected because of the higher dose of paclitaxel that may be administered when given as the 

Genexol-PM formulation. 

NK105 is a micelle formulation of paclitaxel wherein the micelles have been shown to perform 

as true carriers. The NK105 micelles are formed from copolymers with PEG as the hydrophilic 

block and modified PAsp as the core-forming block. The PAsp block is conjugated with 4-phenyl- 

1-butanol as a means to increase the compatibility between the micelle core and paclitaxel. Paclitaxel 

administered in the NK105 formulation has been shown to have a much longer circulation 

lifetime than paclitaxel administered as Taxol w . Specifically, the AUC for paclitaxel was increased 

by 50–86-fold and the t 1/2 was increased 4–6 times for the NK105 formulation in comparison to 

paclitaxel administered as Taxol w . 99 In addition, the therapeutic efficacy of NK105 and Taxol w 

formulations were investigated in an HT-29 colon cancer xenograft model in nude mice. The 

NK105 formulation was found to exhibit statistically superior anti-tumor activity in comparison 

to Taxol w . The anti-tumor activity of NK105 administered at a paclitaxel-equivalent dose of 25 mg/ 

kg was comparable to that obtained following the administration of Taxol w at a paclitaxel dose of 

100 mg/kg. In addition, all tumors disappeared in the NK105 treatment group following one dose of 

100 mg/kg, and all mice remained tumor-free for the 33-day duration of the study. This significant 

improvement in the anti-cancer activity of paclitaxel following encapsulation in this micelle system 

may be attributed to the ability of the NK105 micelles to function as true drug carriers. 

17.6 EVOLUTION OF POLYMERIC MICELLES 

Recently, much effort has been devoted to increasing drug accumulation at diseased sites. Advances 

in synthetic chemistry and a more comprehensive understanding of biological systems have facilitated 

the development of smart micelle systems that are capable of targeting drugs to specific tissues 

or cellular organelles. Through chemical modification, conventional micelles may be designed to 

deliver to and/or release drugs at a specific site of interest by active targeting or 

external stimulation. 

17.6.1 TRIGGERED STIMULATION OF DRUG ACTIVATION OR RELEASE 

Following administration of drug-loaded micelles, external stimuli may be locally applied at the 

tumor in order to release or activate the drug. This type of drug delivery mechanism may ensure a 

more accurate temporal and spatial accumulation of drugs in the tumor as well as reduce toxicity to 

other tissues. In this way, localized drug delivery may be achieved in a non-invasive manner using 

external stimuli such as ultrasound, heat, and light. 

17.6.1.1 Ultrasonic Irradiation 


The application of ultrasonic irradiation at tumor sites has been found to be effective in enhancing 

the accumulation of drugs in tumor cells and has resulted in improved therapeutic efficacy. The use 

of ultrasonic irradiation for micelle-based chemotherapy can be optimized by varying many factors 

including the sonication frequency, the types of waves applied, the time between drug injection and 

the application of ultrasound as well as the physico-chemical properties of the micelles. 150,151 

Micelle properties such as the hydrophobicity and state of the micelle core were found to be the 

most influential in increasing drug uptake in tumor cells upon ultrasonic irradiation. 79,128,152 

To date, several mechanisms have been proposed for enhancing tumor cell uptake of micelleformulated 

drugs via ultrasonic irradiation. It had been suggested that ultrasound promotes extravasation 

of drug-loaded carriers because of increased vasculature permeability. 79 However, a recent 

in vivo study revealed that the application of ultrasound does not enhance extravasation of micelles. 

Rather, the enhanced drug accumulation in tumor cells requires initial accumulation of drug-loaded 



micelles in the tumor interstitium via the enhanced permeability and retention (EPR) effect. 150 This 

indicates that the ultrasonic effect occurs at the cellular level and could potentially play a role in 

micelle–cell interactions. In vitro studies have suggested that ultrasound does not increase the 

extent of endocytosis of intact micelles. 119 In fact, ultrasound has been shown to trigger drug 

release as well as enhance drug uptake by increasing cell membrane permeability. 119,151,152 

The application of ultrasound perturbs the drug-loaded micelles, shifting the equilibrium from 

micelle-encapsulated drug to free drug. In this way, the release of drugs from the micelles is 

triggered. 151,152 Furthermore, ultrasonication can cause reversible cell membrane disruption, 

creating temporary pores in the membrane, resulting in enhanced drug diffusion into the cells 

following drug release. 119 This mechanism is also effective in treating multi-drug resistant 

cancer cells because of the enhanced partitioning of the released drug into the intracellular compartment. 

79,118 The percentage of drug released after ultrasonication was found to decrease with 

increasing hydrophobic interactions between the drug and the micelle core. 151 The use of Pluronic w 

copolymers to form micelles that have liquid-like cores facilitates drug release after ultrasonication. 

Nonetheless, following injection, these micelles are very unstable in vivo and may result in reduced 

drug accumulation at tumor sites. PEG-conjugated distearoylphosphatidylethanolamine (DSPE) 

has been added to Pluronic w micelles to improve micelle stability and was proven to be as effective 

as PEG-b-PBLA micelles that have solid-like cores in ultrasound-enhanced chemotherapy. 150 

Interestingly, after inactivation of ultrasonication, the released drug is re-encapsulated in the 

micelles, and the intact drug-loaded micelles are capable of recirculating until they reach the 

tumor again. As a result, more than a 3-fold increase in drug uptake in tumor cells is observed 

in vivo, and a reduction in the accumulation of the drug in normal tissues is achieved when 

compared to non-sonicated tumors. 150 

17.6.1.2 Thermosensitive Systems 

In addition to ultrasonication, micelles can also be designed to trigger the release of drugs at the 

tumors using heat as an external stimulus. Application of thermo-responsive micelles for the 

treatment of solid tumors involves the use of a thermosensitive polymer as the micelle-building 

block and the local administration of heat at the tumor site. Thermo-responsive copolymers 

including poly(N-isopropylacrylamide-b-D,L-lactide) (PIPAAm-b-PDLLA), poly(N-isopropylacrylamide-b-(butyl 

methacrylate)) (PIPAAm-b-PBMA), poly(N-isopropylacrylamide-co-N,Ndimethylacrylamide)-b-poly(D,L-lactide-co-glycolide) 

(P(IPAAm-co-DMAAm)-b-PLGA) have 

been used for the preparation of thermo-responsive micelles. 122,153–155 These micelles undergo 

reversible structural changes that facilitate drug release above the lower critical solution temperature 

(LCST) of the thermosensitive polymer. In the case of PIPAAm micelles, dehydration of the 

outer shells of the micelles also occurs above the LCST, resulting in the enhanced adsorption of 

drug-loaded micelles to cells via increased hydrophobic interactions. 155–157 

Okano’s group developed a thermo-responsive, doxorubicin-loaded micelle system with an 

LCST of 328C. By toggling the temperature below and above the LCST, the release of doxorubicin 

may be switched reversibly between an on and off state. At temperatures above the LCST, doxorubicin 

release is accelerated, whereas decreasing the temperature to levels below the LCST results 

in minimal drug release. 155 This effect is also influenced by the state of the micelle core in addition 

to the thermosensitivity of the PIPAAm outer shells. For instance, micelles that are formed from 

copolymers with liquid-like cores (e.g., TgZ208C for PBMA) are more sensitive to thermal 

triggered drug release than micelles with solid-like cores (e.g., TgZ1008C for polystyrene (PS)) 

at temperatures above the LCST. As a result, the doxorubicin-loaded PPIAAm-b-PBMA micelles 

were three times more cytotoxic than the PPIAAm-b-PS micelles above the LCST mainly as a 

result of their distinct doxorubicin release profiles. At temperatures below the LCST, both systems 

exhibited almost no cytotoxicity. 153 


342 

For in vivo delivery, the LCST of the thermo-responsive micelles should be several degrees 

above physiological temperature (i.e., LCSTO378C). 157 The LCST can be manipulated by finetuning 

the hydrophilicity of the PIPAAm segment. The addition of hydrophilic groups to the 

PIPAAm polymer can elevate the LCST of the micelles formed. 122,155–158 For example, Liu et 

al. reported that by adding hydrophilic DMAAm to the PIPAAm segment, the P(IPAAm-co- 

DMAAm)-b-PLGA micelles have an LCST of approximately 398C. As a result, the doxorubicin-loaded 

P(IPAAm-co-DMAAm)-b-PLGA micelles have been shown to be more potent in the 4T1 

mouse breast cancer cells at 39.58C (IC50Z3.1 mg/L) than at 378C (IC50Z7.9 mg/L). 122 Therefore, 

heat-activated drug release in tumors (above 378C) can reduce the possible toxicity associated 

with the drug in healthy tissues (at 378C). 

17.6.1.3 Photodynamic Therapy 

In photodynamic therapy (PDT), light is used to activate the drug (a photosensitizer) at diseased 

sites. Initially, PDT was used for the treatment of superficial tumors such as skin cancer. Recent 

improvements in the technology have allowed for the treatment of deep tumors such as lung 

cancers. 159 The tumor area is exposed to light of a specific wavelength that activates the drug, 

leading to the destruction of tumor cells and proximal tumor vasculature via the production of 

reactive oxygen species or by other mechanisms. 160,161 In order for the PDT agents to be effective 

in destroying tumor cells, it is necessary that there is no aggregation of the drug. However, in 

aqueous media, the hydrophobic photosensitizers tend to aggregate, resulting in reduced potency. 

Pluronic w micelles were reported to be able to solubilize benzoporphyrin while avoiding drug 

aggregation in the micelle core, implying that the potency of the photosensitizer was maintained. 162 

In this way, polymeric micelles become an attractive alternative for the solubilization and delivery 

of photosensitizers. 

Kataoka’s group has developed polyion complex (PIC) micelles based on anionic copolymers 

(PEG-b-poly(L-lysine) or PEG-b-PAsp) for the encapsulation of a polycationic photosensitizer. 

These micelles are stabilized by electrostatic interactions, and the micelles are pH-responsive, 

allowing intratumoral drug release or intracellular endosomal escape. 160,162,163 The photosensitizer-loaded 

PIC micelles showed no skin phototoxicity in rats upon exposure to broadband visible 

light, and severe skin phototoxicity was observed with the administration of Photofrin w (a currently 

approved PDT drug formulated in 0.9% sodium chloride). 164 The photosensitizer-loaded PIC 

micelles were also shown to exhibit enhanced efficacy and reduced toxicity in the Lewis lung 

carcinoma (LLC) cell line in comparison to free photosensitizer. 160 

It has also been reported that the subcellular localization of the photosensitizer influences its 

efficacy. Leroux’s group has developed pH-sensitive aluminum chloride pthalocyanine (AICIPc)loaded 

N-isopropylacrylamide copolymer micelles for treatment against EMT-6 tumors in vivo. 

The tumor accumulation of the micelle-formulated drug is lower than that of AICIPc dissolved in 

Cremophor EL. However, the copolymer micelle formulation has been shown to have enhanced 

therapeutic efficacy that is most likely attributed to an increase in the cytoplasmic accumulation of 

the drug. 165,166 In many cases, the accumulation of photosensitizer at the mitochondria has been 

found to be the most efficacious. 161,167 Therefore, the use of pH-responsive polymeric micelles for 

PDT could enhance drug accumulation in the mitochondria following endosomal/lysosomal escape 

of the micelles. In the future, the design of polymeric micelles for use in PDT will likely be aimed at 

improving tumor and subcellular localization by fine-tuning the copolymer composition in order to 

achieve active targeting. 

17.6.2 ACTIVE TARGETING 


The phenomenon of active targeting of micellar delivery vehicles loaded with drugs implies that the 

micelles can accumulate at the targeted cells or tissues without external stimuli via specific sensing 



FIGURE 17.7 (See color insert following page 522.) Confocal fluorescence microscopy image of epidermal 

growth factor receptor (EGFR)-overexpressed breast cancer cells (MDA-MB-468) following incubation with 

EGF-conjugated poly(ethylene glycol)-block-poly(d-valerolactone) micelles (as shown in red) for 2 h at 378C. 

The cell nuclei are stained with Hoechst 33258 (blue). The EGF-conjugated micelles mainly localized at the 

perinuclear region as well as within the nucleus as shown by the white arrows. 

mechanisms that are integrated into the carriers. The carriers can be engineered by means of ligand 

coupling or addition of pH-sensitive moieties according to the biological characteristics of the 

diseased site. In particular, many cancers are associated with over expression of certain cell surface 

receptors and acidosis at the tumor sites. 

17.6.2.1 Ligand Coupling 

It has been demonstrated that ligand-conjugated drug carriers can selectively and efficiently 

accumulate in target cells. 168–172 The addition of biomolecules to the surface of micelles enables 

the drug-loaded micelles to enter the cells via receptor-mediated endocytosis as shown in 

Figure 17.5. Through the highly specific binding affinity between the ligand–receptor complex, 

the attachment of ligands to micelles surface allows for targeted drug delivery to specific cells that 

express the corresponding cell surface receptors. In particular, many types of cancer cells express a 

higher-than-normal level of cell surface receptors that have binding affinities for growth-dependent 

molecules as a result of their rapid growth rate. 173 Ligands such as sugar moieties, 174–177 transferrin, 

178–180 folate, 78,95,181 epidermal growth factor (EGF), 115,167,178 and antibody fragments 182 

have been attached to the surface of copolymer micelles to target different types of cancers. 

Kabanov’s group was one of the first to apply the ligand coupling concept to micellar drug 

delivery systems. The authors demonstrated that by chemically conjugating a2-glycoprotein, a 

ligand that targets brain glial cells, to FITC-containing Pluronic w micelles, the accumulation of 

FITC in brain tissues was increased and clearance of FITC by the lung was decreased in comparison 

to delivery using conventional Pluronic w micelles. 126,183 Recently, another ligand, folate, was 

conjugated to PEG-b-PLGA micelle system for targeting folate receptor over expressing 

cancers. The doxorubicin-loaded folate-PEG-b-PLGA micelles were found to enhance the intracellular 

accumulation of doxorubicin and resulted in a lower IC50 in folate receptor overexpressing 

KB epidermal carcinoma cells. In vivo studies revealed that treatment with doxorubicin-loaded 

folate-PEG-b-PLGA micelles resulted in a high suppression of tumor growth. 78 


344 

In cases where nuclear targeting is desired, ligands such as EGF and the HIV-1 TAT peptide 

that have nuclear translocation properties, may be employed. 6,115,184 Zeng et al. reported on an 

EGF-conjugated PEG-b-poly(d-valerolactone) (PEG-b-PVL) micelle system that targets EGFRover 

expressing breast cancers. The EGF-PEG-b-PVL micelles were shown to mainly localize in 

the perinuclear region and in the nucleus of MDA-MB-468 breast cancer cells as shown in 

Figure 17.7. 115 Nuclear targeting is critical for the delivery of anti-cancer drugs and oligonucleotides 

whose site of action is located in the nucleus. Wagner’s group has developed EGF- and 

transferrin-conjugated PEG-b-poly(ethylenimine) (PEG-b-PEI) micelle systems for gene delivery 

to tumor tissues. 178,180,185,186 Molecules such as oligonucleotides are more susceptible to 

degradation in the endosomal or lysosomal compartments; therefore, pH-responsive micelles 

that are capable of destabilizing the endosomal or lysosomal membranes could potentially 

improve the therapeutic effect of these molecules. 

17.6.2.2 pH-Responsive Systems 


The pH where drug release is triggered can be adjusted by varying the pH-sensitive moiety 

incorporated in the copolymer. 158,187 In this way, the pH phase transition of a micelle system 

can be designed to suit targets that have different pH values. For instance, micelles that are 

stable at pHw2 and are disrupted at pHO7.2 may be suitable for oral delivery of hydrophobic 

drugs. As an example, a 10-fold increase in the bioavailability of fenofibrate formulated in pH-responsive 

micelles was reported. 188 In the case of tumor targeting, micelles with a phase transition of 

approximately pH 7.0 could be beneficial as tumor tissues are known to be slightly more acidic 

(mean pH 7.0) than normal tissues (pH 7.4). 187,189,190 If subcellular targeting is desired, cytoplasmic 

accumulation of drugs can be enhanced by utilizing micelles that are responsive to pH values 

ranging from 5.0 to 6.0 (endosomal/lysosomal pH). 187,191,192 

pH-responsive micelles have also been developed by conjugating drugs to the hydrophobic 

blocks of copolymers using pH-sensitive linkages. For instance, doxorubicin can be conjugated to 

PEG-b-P(Asp) copolymers via a hydrazone linkage. The micelles formed from this copolymer were 

found to be extracellularly stable, and they provided selective release of doxorubicin in the endosomes 

where the pH ranges from 5.0 to 6.0. The released drug was found to first localize in the 

cytoplasm and eventually in the nucleus. The pH-responsive doxorubicin micelles were also shown 

to be capable of infiltrating avascular multicellular tumor spheroids in a uniform manner. Furthermore, 

the micelles were effective in suppressing tumor growth in a murine C26 colon tumor model 

and increased the MTD by more than 2.5 times because of their secure drug encapsulation. 191,192 

Another approach to develop pH-responsive micelles is to disrupt micelle formation in the 

acidic environment by introducing titrable groups into the copolymer. 187 Leroux’s group introduced 

a hydrophilic titrable monomer (methacrylic acid or MAA) into the poly(Nisopropylacrylamide)-co-octadecyl 

acrylate (poly(NIPA)-co-ODA) copolymer to form pH-responsive 

micelles for PDT. 158 These micelles were found to undergo a pH phase transition and structural 

changes in the micelle core at a pH of approximately 5.8. The enhanced cytotoxicity of the AICIPcloaded 

poly(NIPA)-co-ODA-co-MAA micelles were found to be highly dependent on the pH-gradient 

between the cytoplasm and the endosomal/lysosomal compartments. 158 

Bae and co-workers have developed several pH-responsive micelle systems by using poly 

(L-histidine) (polyHis) as the pH-sensitive moiety. The CMC of the PEG-b-polyHis copolymer 

was found to increase as the pH of the media was decreased with an abrupt increase in the CMC 

occurring at a pH of 7.2. Therefore, with the appropriate alteration in pH, these micelles become 

thermodynamically unstable and dissociate, resulting in the release of their contents. 193 Doxorubicin-loaded 

PEG-b-polyHis micelles were shown to result in a significant increase in the in vitro 

intracellular accumulation of drug, in vivo tumor localization of drug, and tumor growth inhibition 

in the A2780 ovarian tumor model via extracellular pH-dependent targeting of the tumor. 194 The 

addition of polyHis to PEG-b-PDLLA further reduces the pH phase transition of the micelles to the 



range of 6.6–7.2. These micelles were further modified with ligand coupling to achieve a doubletargeting 

effect. 181,190 For example, Lee et al. prepared polyHis-b-PEG-b-PDLLA mixed micelles 

with biotin as the targeting ligand. Biotin is protected by the PEG shell at pHO7.0 and is exposed 

on the surface of micelles when the pH is decreased to 6.5–7.0 (extracellular tumor pH). As the pH 

is further reduced to less than 6.5 (endosomal pH), the micelles destabilize, triggering drug release 

and disrupting endosomal membranes via the proton-sponge effect. 190 In the future, it is anticipated 

that combinations of sensing mechanisms will be used to improve the effectiveness of micellar-drug 

targeting to the site of interest. 

17.7 FUTURE PERSPECTIVES 

The field of block copolymer micelles for drug delivery is rapidly expanding and has great potential 

in cancer therapeutics. Although there are four promising polymeric micelle formulations currently 

in clinical trial development, many preclinical fundamental studies are on-going in order to fully 

exploit the use of block copolymer micelles for drug delivery. The advances in synthetic polymer 

science continue to provide new biodegradable and biocompatible polymers that can be tailored and 

designed for the challenges that anti-cancer drug delivery continues to face. By enhancing the 

compatibility between the copolymer and the drug, block copolymer micelles have emerged as true 

carriers rather than solubilizers for pharmaceutical agents. The mechanisms of loading and release of 

drugs from micelles are being elucidated and will provide a greater understanding of the 

relationships between the physico-chemical properties of micelles and their invivo performance. 

In the future, multiple targeting mechanisms may be integrated into one micelle system to fully 

exploit the use of polymeric micelle as delivery vehicles for anti-cancer drugs. 


The authors are grateful to NSERC for funding their research in the area of block copolymer 

micelles for applications in drug delivery. In addition, the authors thank Lorne F. Lambier for a 

scholarship to H. Lee, CIHR/Rx&D for a career award to C. Allen, a post-doctoral fellowship to P. 

Lim Soo, NSERC for a scholarship to J. Liu, and OGS for a scholarship to M. Butler. 

REFERENCES 



2. Torchilin, V. P., Recent advances with liposomes as pharmaceutical carriers, Nature Reviews Drug 

Discovery, 4, 145–160, 2005. 

3. Allen, C., Maysinger, D., and Eisenberg, A., Nano-engineering block copolymer aggregates for drug 

delivery, Colloids and Surfaces B, 16, 3–27, 1999. 

4. Kwon, G. S., Diblock copolymer nanoparticles for drug delivery, Critical Reviews in Therapeutic 

Drug Carrier Systems, 15, 481–512, 1998. 

5. Kataoka, K., Kwon, G. S., Yokoyama, M., Okano, T., and Sakurai, Y., Block-copolymer micelles as 

vehicles for drug delivery, Journal of Controlled Release, 24, 119–132, 1993. 

6. Gaucher, G., Dufresne, M. H., Sant, V. P., Kang, N., Maysinger, D., and Leroux, J. C., Block 

copolymer micelles: Preparation, characterization and application in drug delivery, Journal of 


7. Kabanov, A. V., Batrakova, E. V., and Alakhov, V. Y., Pluronic (R) block copolymers as novel 

polymer therapeutics for drug and gene delivery, Journal of Controlled Release, 82, 189–212, 2002. 

8. Elbert, D. L. and Hubbell, J. A., Surface treatments of polymers for biocompatibility, Annual Review 

of Material Science, 26, 365–394, 1996. 


346 


9. Harris, J. M., Ed., Poly(ethylene glycol) chemistry: Biotechnical and biomedical applications, In 

Topics in Applied Chemistry, Plenum, New York; London, 385, 1992. 

10. Allen, C., Dos Santos, N., Gallagher, R., Chiu, G. N. C., Shu, Y., Li, W. M., Johnstone, S. A., Janoff, 

A. S., Mayer, L. D., Webb, M. S., and Bally, M. B., Controlling the physical behavior and biological 

performance of liposome formulations through use of surface grafted poly(ethylene glycol), 

Bioscience Representatives, 22, 225–250, 2002. 

11. Gref, R., Domb, A., Quellec, P., Blunk, T., Muller, R. H., Verbavatz, J. M., and Langer, R., The 

controlled intravenous delivery of drugs using peg-coated sterically stabilized nanospheres, 


12. Hunter, R. J. and White, L. R., Foundations of colloid science, Oxford Science Publications, Clarendon 

Press, Oxford [Oxfordshire]; Oxford University Press, New York, p. 2 v. (1089), 1987. 

13. Müller, R. H., Colloidal Carriers for Controlled Drug Delivery and Targeting: Modification, 

Characterization and In Vivo Distribution, Wissenschaftliche Verlagsgesellschaft, Stuttgart; CRC 

Press; Sole distributor for North America CRC Press Inc., Boca Raton, FL, p. 379, 1991. 

14. Kato, Y., Watanabe, K., Nakakura, M., Hosokawa, T., Hayakawa, E., and Ito, K., Blood clearance 

and tissue distribution of various formulations of alpha-tocopherol injection after intravenous administration, 

Chemical and Pharmaceutical Bulletin (Tokyo), 41, 599–604, 1993. 

15. Oku, N. and Namba, Y., Long-circulating liposomes, Critical Reviews in Therapeutic Drug Carrier 

Systems, 11, 231–270, 1994. 

16. Nakanishi, T., Fukushima, S., Okamoto, K., Suzuki, M., Matsumura, Y., Yokoyama, M., Okano, T., 

Sakurai, Y., and Kataoka, K., Development of the polymer micelle carrier system for doxorubicin, 


17. Seymour, L. W., Duncan, R., Strohalm, J., and Kopecek, J., Effect of molecular weight (Mw) ofN- 

(2-hydroxypropyl)methacrylamide copolymers on body distribution and rate of excretion after 

subcutaneous, intraperitoneal, and intravenous administration to rats, Journal of Biomedical 

Materials Research, 21, 1341–1358, 1987. 

18. Allen, C., Yu, Y., Maysinger, D., and Eisenberg, A., Polycaprolactone-b-poly(ethylene oxide) block 

copolymer micelles as a novel drug delivery vehicle for neurotrophic agents FK506 and L-685,818, 


19. Zhang, L. F. and Eisenberg, A., Formation of crew-cut aggregates of various morphologies 

from amphiphilic block copolymers in solution, Polymers for Advanced Technologies, 9, 

677–699, 1998. 

20. Choucair, A. and Eisenberg, A., Control of amphiphilic block copolymer morphologies using 

solution conditions, European Physical Journal E, 10, 37–44, 2003. 

21. Discher, D. E. and Eisenberg, A., Polymer vesicles, Science, 297, 967–973, 2002. 

22. Lim Soo, P. and Eisenberg, A., Preparation of block copolymer vesicles in solution, Journal of 

Polymer Science Part B-Polymer Physics, 42, 923–938, 2004. 

23. Kabanov, A. V., Nazarova, I. R., Astafieva, I. V., Batrakova, E. V., Alakhov, V. Y., Yaroslavov, 

A. A., and Kabanov, V. A., Micelle formation and solubilization of fluorescent-probes in poly(oxyethylene-b-oxypropylene-b-oxyethylene) 


24. Wilhelm, M., Zhao, C. L., Wang, Y. C., Xu, R. L., Winnik, M. A., Mura, J. L., Riess, G., and 

Croucher, M. D., Polymer micelle formation. 3. Poly(styrene-ethylene oxide) block copolymer 

micelle formation in water—a fluorescence probe study, Macromolecules, 24, 1033–1040, 1991. 

25. Liu, T. B., Kim, K., Hsiao, B. S., and Chu, B., Regular and irregular micelles formed by A LEL 

triblock copolymer in aqueous solution, Polymer, 45, 7989–7993, 2004. 

26. Alakhov, V., Klinski, E., Li, S. M., Pietrzynski, G., Venne, A., Batrakova, E., Bronitch, T., and 

Kabanov, A., Block copolymer-based formulation of doxorubicin. From cell screen to clinical trials, 

Colloids and Surfaces B: Biointerfaces, 16, 113–134, 1999. 

27. Lee, H., Zeng, F. Q., Dunne, M., and Allen, C., Methoxy poly(ethylene glycol)-block-poly(deltavalerolactone) 

copolymer micelles for formulation of hydrophobic drugs, Biomacromolecules, 6, 

3119–3128, 2005. 

28. Kang, N. and Leroux, J. C., Triblock and star-block copolymers of N-(2-hydroxypropyl) methacrylamide 

or N-vinyl-2-pyrrolidone and D,L-lactide: Synthesis and self-assembling properties in water, 

Polymer, 45, 8967–8980, 2004. 



29. Alexandridis, P., Holzwarth, J. F., and Hatton, T. A., Micellization of poly(ethylene oxide)–poly(propylene 

oxide)–poly(ethylene oxide) triblock copolymers in aqueous-solutions—thermodynamics 

of copolymer association, Macromolecules, 27, 2414–2425, 1994. 

30. Yamamoto, Y., Yasugi, K., Harada, A., Nagasaki, Y., and Kataoka, K., Temperature-related change 

in the properties relevant to drug delivery of poly(ethylene glycol)–poly(D,L-lactide) block copolymer 

micelles in aqueous milieu, Journal of Controlled Release, 82, 359–371, 2002. 

31. Kang, N., Perron, M. E., Prud’homme, R. E., Zhang, Y. B., Gaucher, G., and Leroux, J. C., Stereocomplex 

block copolymer micelles: Core-shell nanostructures with enhanced stability, Nano Letters, 

5, 315–319, 2005. 

32. Yokoyama, M., Fukushima, S., Uehara, R., Okamoto, K., Kataoka, K., Sakurai, Y., and Okano, T., 

Characterization of physical entrapment and chemical conjugation of adriamycin in polymeric 

micelles and their design for in vivo delivery to a solid tumor, Journal of Controlled Release, 50, 

79–92, 1998. 

33. Croy, S. R. and Kwon, G. S., The effects of pluronic block copolymers on the aggregation state of 

nystatin, Journal of Controlled Release, 95, 161–171, 2004. 

34. Batrakova, E., Lee, S., Li, S., Venne, A., Alakhov, V., and Kabanov, A., Fundamental relationships 

between the composition of pluronic block copolymers and their hypersensitization effect in MDR 

cancer cells, Pharmaceutical Research, 16, 1373–1379, 1999. 

35. Burt, H. M., Zhang, X. C., Toleikis, P., Embree, L., and Hunter, W. L., Development of copolymers 

of poly(D,L-lactide) and methoxypolyethylene glycol as micellar carriers of paclitaxel, Colloids and 

Surfaces B: Biointerfaces, 16, 161–171, 1999. 

36. Liu, J., Xiao, Y., and Allen, C., Polymer–drug compatibility: A guide to the development of delivery 

systems for the anticancer agent, ellipticine, Journal of Pharmaceutical Sciences, 93, 132–143, 

2004. 

37. Zhang, X. C., Jackson, J. K., and Burt, H. M., Development of amphiphilic diblock copolymers as 

micellar carriers of taxol, International Journal of Pharmaceutics, 132, 195–206, 1996. 

38. Yokoyama, M., Opanasopit, P., Okano, T., Kawano, K., and Maitani, Y., Polymer design and 

incorporation methods for polymeric micelle carrier system containing water-insoluble anticancer 

agent camptothecin, Journal of Drug Targeting, 12, 373–384, 2004. 

39. Kohori, F., Yokoyama, M., Sakai, K., and Okano, T., Process design for efficient and controlled drug 

incorporation into polymeric micelle carrier systems, Journal of Controlled Release, 78, 155–163, 

2002. 

40. Jette, K. K., Law, D., Schmitt, E. A., and Kwon, G. S., Preparation and drug loading of poly(ethylene 

glycol)-block-poly(epsilon-caprolactone) micelles through the evaporation of a cosolvent azeotrope, 

Pharmaceutical Research, 21, 1184–1191, 2004. 

41. Shuai, X., Ai, H., Nasongkla, N., Kim, S., and Gao, J., Micellar carriers based on block copolymers 

of poly(epsilon-caprolactone) and poly(ethylene glycol) for doxorubicin delivery, Journal of 


42. Zhao, J. X., Allen, C., and Eisenberg, A., Partitioning of pyrene between “crew cut” block copolymer 

micelles and H 2O/DMF solvent mixtures, Macromolecules, 30, 7143–7150, 1997. 

43. Sharma, P. K. and Bhatia, S. R., Effect of anti-inflammatories on pluronic (R) F127: Micellar 

assembly, gelation and partitioning, International Journal of Pharmaceutics, 278, 361–377, 2004. 

44. Barreiro-Iglesias, R., Bromberg, L., Temchenko, M., Hatton, T. A., Concheiro, A., and Alvarez- 

Lorenzo, C., Solubilization and stabilization of camptothecin in micellar solutions of pluronic-gpoly(acrylic 

acid) copolymers, Journal of Controlled Release, 97, 537–549, 2004. 

45. Lim Soo, P., Luo, L., Maysinger, D., and Eisenberg, A., Incorporation and release of hydrophobic 

probes in biocampatible polycaprolactone-block-poly(ethylene oxide) micelles: Implications for 

drug delivery, Langmuir, 18, 9996–10004, 2002. 

46. Letchford, K., Zastre, J., Liggins, R., and Burt, H., Synthesis and micellar characterization of short 

block length methoxy poly(ethylene glycol)-block-poly(caprolactone) diblock copolymers, Colloids 

and Surfaces B: Biointerfaces, 35, 81–91, 2004. 

47. Choucair, A. and Eisenberg, A., Interfacial solubilization of model amphiphilic molecules in block 

copolymer micelles, Journal of the American Chemical Society, 125, 11993–12000, 2003. 


348 


48. Goldenberg, M. S., Bruno, L. A., and Rennwantz, E. L., Determination of solubilization sites and 

efficiency of water-insoluble agents in ethylene oxide-containing nonionic micelles, Journal of 

Colloid and Interface Science, 158, 351–363, 1993. 

49. Nagarajan, R., Solubilization of “guest” molecules into polymeric aggregates, Polymers for 

Advanced Technologies, 12, 23–43, 2001. 

50. Nagarajan, R. and Ganesh, K., Comparison of solubilization of hydrocarbons in (PEO–PPO) diblock 

versus (PEO–PPO–PEO) triblock copolymer micelles, Journal of Colloid and Interface Science, 

184, 489–499, 1996. 

51. Kozlov, M. Y., Melik-Nubarov, N. S., Batrakova, E. V., and Kabanov, A. V., Relationship between 

pluronic block copolymer structure, critical micellization concentration and partitioning coefficients 

of low molecular mass solutes, Macromolecules, 33, 3305–3313, 2000. 

52. Xing, L. and Mattice, W. L., Strong solubilization of small molecules by triblock-copolymer 

micelles in selective solvents, Macromolecules, 30, 1711–1717, 1997. 

53. Teng, Y., Morrison, M. E., Munk, P., Webber, S. E., and Prochazka, K., Release kinetics studies of 

aromatic molecules into water from block polymer micelles, Macromolecules, 31, 3578–3587, 1998. 

54. Lin, W. J., Juang, L. W., and Lin, C. C., Stability and release performance of a series of pegylated 

copolymeric micelles, Pharmaceutical Research, 20, 668–673, 2003. 

55. Gadelle, F., Koros, W. J., and Schechter, R. S., Solubilization of aromatic solutes in block copolymers, 

Macromolecules, 28, 4883–4892, 1995. 

56. Lee, J., Cho, E. C., and Cho, K., Incorporation and release behavior of hydrophobic drug in functionalized 

poly(D,L-lactide)-block-poly(ethylene oxide) micelles, Journal of Controlled Release, 94, 

323–335, 2004. 

57. Lavasanifar, A., Samuel, J., and Kwon, G. S., Micelles self-assembled from poly(ethylene oxide)block-poly(N-hexyl 

stearate L-aspartamide) by a solvent evaporation method: Effect on the solubilization 

and haemolytic activity of amphotericin B, Journal of Controlled Release, 77, 155–160, 

2001. 

58. Adams, M. L. and Kwon, G. S., Relative aggregation state and hemolytic activity of amphotericin B 

encapsulated by poly(ethylene oxide)-block-poly(N-hexyl-L-aspartamide)-acyl conjugate micelles: 

Effects of acyl chain length, Journal of Controlled Release, 87, 23–32, 2003. 

59. Oh, K. T., Bronich, T. K., and Kabanov, A. V., Micellar formulations for drug delivery based on 

mixtures of hydrophobic and hydrophilic pluronic((R)) block copolymers, Journal of Controlled 

Release, 94, 411–422, 2004. 

60. Yokoyama, M., Satoh, A., Sakurai, Y., Okano, T., Matsumura, Y., Kakizoe, T., and Kataoka, K., 

Incorporation of water-insoluble anticancer drug into polymeric micelles and control of their particle 

size, Journal of Controlled Release, 55, 219–229, 1998. 

61. Bromberg, L. and Magner, E., Release of hydrophobic compounds from micellar solutions of 

hydrophobically modified polyelectrolytes, Langmuir, 15, 6792–6798, 1999. 

62. Kim, S. Y., Ha, J. C., and Lee, Y. M., Poly(ethylene oxide)–poly(propylene oxide)–poly(ethylene 

oxide)/poly(epsilon-caprolactone) (PCL) amphiphilic block copolymeric nanospheres. II. Thermoresponsive 

drug release behaviors, Journal of Controlled Release, 65, 345–358, 2000. 

63. Kim, S. Y., Kim, J. H., Kim, D., An, J. H., Lee, D. S., and Kim, S. C., Drug-releasing kinetics of 

MPEG/PLLA block copolymer micelles with different PLLA block lengths, Journal of Applied 

Polymer Science, 82, 2599–2605, 2001. 

64. Cho, Y. W., Lee, J., Lee, S. C., Huh, K. M., and Park, K., Hydrotropic agents for study of in vitro 

paclitaxel release from polymeric micelles, Journal of Controlled Release, 97, 249–257, 2004. 

65. Lim Soo, P., Lovric, J., Davidson, P., Maysinger, D., and Eisenberg, A., Polycaprolactone-blockpoly(ethylene 

oxide) micelles: A nanodelivery system for 17beta-estradiol, Molecular Pharmacology, 

2, 519–527, 2005. 


long-circulating polymeric nanospheres, Science, 263, 1600–1603, 1994. 

67. Nah, J. W., Jeong, Y. I., and Cho, C. S., Norfloxacin release from polymeric micelle of poly(gammabenzyl 

L-glutamate) poly(ethylene oxide) poly(gamma-benzylL-glutamate) block copolymer, 

Bulletin of the Korean Chemical Society, 19, 962–967, 1998. 



68. Ryu, J., Jeong, Y. I., Kim, I. S., Lee, J. H., Nah, J. W., and Kim, S. H., Clonazepam release from coreshell 

type nanoparticles of poly(epsilon-caprolactone)/poly(ethylene glycol)/poly(epsilon-caprolactone) 

triblock copolymers, International Journal of Pharmaceutics, 200, 231–242, 2000. 

69. Kim, S. Y., Shin, I. L. G., Lee, Y. M., Cho, C. S., and Sung, Y. K., Methoxy poly(ethylene glycol) 

and epsilon-caprolactone amphiphilic block copolymeric micelle containing indomethacin. II. 

Micelle formation and drug release behaviours, Journal of Controlled Release, 51, 13–22, 1998. 

70. Nah, J. W., Jeong, Y. I., Cho, C. S., and Kim, S. I., Drug-delivery system based on core-shell-type 

nanoparticles composed of poly(gamma-benzyl-L-glutamate) and poly(ethylene oxide), Journal of 

Applied Polymer Science, 75, 1115–1126, 2000. 

71. Leroux, J. C., Allemann, E., DeJaeghere, F., Doelker, E., and Gurny, R., Biodegradable nanoparticles—from 

sustained release formulations to improved site specific drug delivery, Journal of 


72. Kwon, G. S., Naito, M., Kataoka, K., Yokoyama, M., Sakurai, Y., and Okano, T., Block copolymer 

micelles as vehicles for hydrophobic drugs, Colloids and Surfaces B: Biointerfaces, 2, 429–434, 

1994. 

73. Saltzman, W. M., Drug delivery: Engineering principles for drug delivery, Topics in Chemical 

Engineering, Oxford University Press, New York, 372, 2001. 

74. Doxorubicin, U.S. National Library of Medicine and National Institutes of Health http://www.nlm. 

nih.gov/medlineplus/druginfo/medmaster/a682221.html. 

75. Lipp, H.-P. and Bokemeyer, C., Anticancer drug toxicity, In Anthracyclines and Other Intercalating 

Agents, Lipp, H. P., Ed., Marcel Dekker Inc., New York, pp. 81–113, 1999. 

76. Kwon, G. S., Yokoyama, M., Okano, T., Sakurai, Y., and Kataoka, K., Biodistribution of micelleforming 

polymer–drug conjugates, Pharmaceutical Research, 10, 970–974, 1993. 

77. Yoo, H. S. and Park, T. G., Biodegradable polymeric micelles composed of doxorubicin conjugated 

PLGA–PEG block copolymer, Journal of Controlled Release, 70, 63–70, 2001. 



79. Rapoport, N., Pitt, W. G., Sun, H., and Nelson, J. L., Drug delivery in polymeric micelles: From 

in vitro to in vivo, Journal of Controlled Release, 91, 85–95, 2003. 

80. Pruitt, J. D. and Pitt, W. G., Sequestration and ultrasound-induced release of doxorubicin from 

stabilized pluronic P105 micelles, Drug Delivery, 9, 253–258, 2002. 

81. Danson, S., Ferry, D., Alakhov, V., Margison, J., Kerr, D., Jowle, D., Brampton, M., Halbert, G., and 

Ranson, M., Phase I dose escalation and pharmacokinetic study of pluronic polymer-bound doxorubicin 

(SP1049C) in patients with advanced cancer, British Journal of Cancer, 90, 2085–2091, 2004. 

82. Matsumura, Y., Hamaguchi, T., Ura, T., Muro, K., Yamada, Y., Shimada, Y., Shirao, K., Okusaka, 

T., Ueno, H., Ikeda, M., and Watanabe, N., Phase I clinical trial and pharmacokinetic evaluation of 

NK911, a micelle-encapsulated doxorubicin, British Journal of Cancer, 91, 1775–1781, 2004. 

83. Kwon, G., Suwa, S., Yokoyama, M., Okano, T., Sakurai, Y., and Kataoka, K., Enhanced tumor 

accumulation and prolonged circulation times of micelle-forming poly(ethylene oxide-aspartate) 

block copolymer–adriamycin conjugates, Journal of Controlled Release, 29, 17–23, 1994. 

84. Dollery, C. T., Therapeutic Drugs, Churchill Livingstone, Edinburgh, 1999. 

85. Yokoyama, M., Miyauchi, M., Yamada, N., Okano, T., Sakurai, Y., Kataoka, K., and Inoue, S., 

Polymer micelles as novel drug carrier—adriamycin-conjugated poly(ethylene glycol) poly(aspartic 

acid) block copolymer, Journal of Controlled Release, 11, 269–278, 1990. 

86. Fukushima, S., Machida, M., Akutsu, T., Skimizu, K., Tanaka, S., Okamoto, K., and Mashiba, H., 

Roles of adriamycin and adriamycin dimer in antitumor activity of the polymeric micelle carrier 

system, Colloids and Surfaces B: Biointerfaces, 16, 227–236, 1999. 

87. Paclitaxel, RxList Inc., 2005 http://www.rxlist.com/cgi/generic/paclitaxel.htm. 

88. Liggins, R. T., Hunter, W. L., and Burt, H. M., Solid-state characterization of paclitaxel, Journal of 

Pharmaceutical Sciences, 86, 1458–1463, 1997. 

89. Estimated by CAChe 6.1 (computer-aided molecular design modeling software). 

90. Kim, T. Y., Kim, D. W., Chung, J. Y., Shin, S. G., Kim, S. C., Heo, D. S., Kim, N. K., and Bang, 

Y. J., Phase I and pharmacokinetic study of genexol-PM, a cremophor-free, polymeric micelleformulated 

paclitaxel, in patients with advanced malignancies, Clinical Cancer Research, 10, 

3708–3716, 2004. 


350 


91. Liggins, R. T. and Burt, H. M., Polyether–polyester diblock copolymers for the preparation of 

paclitaxel loaded polymeric micelle formulations, Advanced Drug Delivery Reviews, 54, 

191–202, 2002. 

92. Le Garrec, D., Gori, S., Luo, L., Lessard, D., Smith, D. C., Yessine, M. A., Ranger, M., and Leroux, 

J. C., Poly(N-vinylpyrrolidone)-block-poly(D,L-lactide) as a new polymeric solubilizer for hydrophobic 

anticancer drugs: In vitro and in vivo evaluation, Journal of Controlled Release, 99, 83–101, 

2004. 

93. Shuai, X., Merdan, T., Schaper, A. K., Xi, F., and Kissel, T., Core-cross-linked polymeric micelles as 

paclitaxel carriers, Bioconjugate Chemistry, 15, 441–448, 2004. 

94. Kim, S. Y. and Lee, Y. M., Taxol-loaded block copolymer nanospheres composed of methoxy 

poly(ethylene glycol) and poly(epsilon-caprolactone) as novel anticancer drug carriers, Biomaterials, 

22, 1697–1704, 2001. 

95. Park, E. K., Lee, S. B., and Lee, Y. M., Preparation and characterization of methoxy poly(ethylene 

glycol)/poly(epsilon-caprolactone) amphiphilic block copolymeric nanospheres for tumor-specific 

folate-mediated targeting of anticancer drugs, Biomaterials, 26, 1053–1061, 2005. 

96. Park, S. R., Oh, D. Y., Kim, D. W., Kim, T. Y., Heo, D. S., Bang, Y. J., Kim, N. K., Kang, W.-K., 

Kim, H.-T., Im, S.-A., Kim, J.-H., and Kim, H.-K., A multi-center, late phase II clinical trial of 

genexol (paclitaxel) and cisplatin for patients with advanced gastric cancer, Oncology Reports, 12, 

1059–1064, 2004. 

97. Kim, S. C., Kim, D. W., Shim, Y. H., Bang, J. S., Oh, H. S., Wan Kim, S., and Seo, M. H., In vivo 

evaluation of polymeric micellar paclitaxel formulation: Toxicity and efficacy, Journal of Controlled 

Release, 72, 191–202, 2001. 

98. Yu, J. J., Jeong, Y. I., Shim, Y. H., and Lim, G. T., Preparation of core-shell type nanoparticles of 

diblock copolymers of poly(L-lactide)/poly(ethylene glycol) and their characterization in vitro, 

Journal of Applied Polymer Science, 85, 2625–2634, 2002. 

99. Hamaguchi, T., Matsumura, Y., Suzuki, M., Shimizu, K., Goda, R., Nakamura, I., Nakatomi, I., 

Yokoyama, M., Kataoka, K., and Kakizoe, T., NK105, a paclitaxel-incorporating micellar nanoparticle 

formulation, can extend in vivo antitumour activity and reduce the neurotoxicity of paclitaxel, 

British Journal of Cancer, 92, 1240–1246, 2005. 

100. Watanabe, M., Kawano, K., Yokoyama, M., Opanasopit, P., Okano, T., and Maitani, Y., Preparation 

of camptothecin-loaded polymeric micelles and evaluation of their incorporation and circulation 

stability, International Journal of Pharmaceutics, 308, 183–189, 2006. 

101. Nishiyama, N. and Kataoka, K., Preparation and characterization of size-controlled polymeric 

micelle containing cis-dichlorodiammineplatinum(II) in the core, Journal of Controlled Release, 

74, 83–94, 2001. 

102. Uchino, H., Matsumura, Y., Negishi, T., Koizumi, F., Hayashi, T., Honda, T., Nishiyama, N., Kataoka, 

K., Naito, S., and Kakizoe, T., Cisplatin-incorporating polymeric micelles (NC-6004) can reduce 

nephrotoxicity and neurotoxicity of cisplatin in rats, British Journal of Cancer, 93, 678–687, 2005. 

103. Maysinger, D., Berezovska, O., Savic, R., Lim Soo, P., and Eisenberg, A., Block copolymers modify 

the internalization of micelle-incorporated probes into neural cells, Biochimica et Biophysica Acta, 

1539, 205–217, 2001. 

104. Allen, C., Yu, Y., Eisenberg, A., and Maysinger, D., Cellular internalization of PCL(20)-b-PEO(44) 

block copolymer micelles, Biochimica et Biophysica Acta, 1421, 32–38, 1999. 

105. Savic, R., Luo, L., Eisenberg, A., and Maysinger, D., Micellar nanocontainers distribute to defined 

cytoplasmic organelles, Science, 300, 615–618, 2003. 

106. Luo, L., Tam, J., Maysinger, D., and Eisenberg, A., Cellular internalization of poly(ethylene oxide)b-poly(epsilon-caprolactone) 

diblock copolymer micelles, Bioconjugate Chemistry, 13, 1259–1265, 

2002. 

107. Zastre, J., Jackson, J., and Burt, H., Evidence for modulation of P-glycoprotein-mediated efflux by 

methoxypolyethylene glycol-block-polycaprolactone amphiphilic diblock copolymers, Pharmaceutical 

Research, 21, 1489–1497, 2004. 

108. Alakhov, V., Moskaleva, E., Batrakova, E. V., and Kabanov, A. V., Hypersensitization of multidrug 

resistant human ovarian carcinoma cells by pluronic P85 block copolymer, Bioconjugate Chemistry, 

7, 209–216, 1996. 



109. Batrakova, E. V., Li, S., Elmquist, W. F., Miller, D. W., Alakhov, V. Y., and Kabanov, A. V., 

Mechanism of sensitization of MDR cancer cells by pluronic block copolymers: Selective energy 

depletion, British Journal of Cancer, 85, 1987–1997, 2001. 

110. Batrakova, E. V., Han, H. Y., Alakhov, V., Miller, D. W., and Kabanov, A. V., Effects of pluronic 

block copolymers on drug absorption in Caco-2 cell monolayers, Pharmaceutical Research, 15, 

850–855, 1998. 

111. Zastre, J., Jackson, J., Bajwa, M., Liggins, R., Iqbal, F., and Burt, H., Enhanced cellular accumulation 

of a P-glycoprotein substrate, rhodamine-123, by Caco-2 cells using low molecular weight 

methoxypolyethylene glycol-block-polycaprolactone diblock copolymers, European Journal of 

Pharmaceutics and Biopharmaceutics, 54, 299–309, 2002. 

112. Robinson, J. R. and Lee, V. H. L., Eds., Controlled drug delivery: Fundamentals and applications, 

Drugs and the Pharmaceutical Sciences, Vol. 29, Marcel Dekker, New York, pp. xix, 716, 1987. 

113. Pastan, I. H. and Willingham, M. C., Eds., Endocytosis, Plenum Press, New York, pp. xviii, 326, 

1985. 

114. Torchilin, V. P., Targeted polymeric micelles for delivery of poorly soluble drugs, Cellular and 

Molecular Life Sciences, 61, 2549–2559, 2004. 

115. Zeng, F., Lee, H., and Allen, C., Epidermal growth factor-conjugated poly(ethylene glycol)-blockpoly(d-valerolactone) 

copolymer micelles for targeted delivery of chemotherapeutics, Bioconjugate 

Chemistry, 2006, in Press. 


compounds, Advanced Drug Delivery Reviews, 54, 203–222, 2002. 

117. Muniruzzaman, M., Marin, A., Luo, Y., Prestwich, G. D., Pitt, W. G., Husseini, G., and Rapoport, 

N. Y., Intracellular uptake of pluronic copolymer: Effects of the aggregation state, Colloids and 


118. Rapoport, N., Marin, A., Luo, Y., Prestwich, G. D., and Muniruzzaman, M. D., Intracellular uptake 

and trafficking of pluronic micelles in drug-sensitive and MDR cells: Effect on the intracellular drug 

localization, Journal of Pharmaceutical Science, 91, 157–170, 2002. 

119. Husseini, G. A., Runyan, C. M., and Pitt, W. G., Investigating the mechanism of acoustically 

activated uptake of drugs from pluronic micelles, BMC Cancer, 2, 20, 2002. 

120. Liu, J., Zeng, F., and Allen, C., Influence of serum protein on polycarbonate-based copolymer 

micelles as a delivery system for a hydrophobic anti-cancer agent, Journal of Controlled Release, 

103, 481–497, 2005. 

121. Pante, N. and Kann, M., Nuclear pore complex is able to transport macromolecules with diameters of 

about 39 nm, Molecular Biology of the Cell, 13, 425–434, 2002. 

122. Liu, S. Q., Tong, Y. W., and Yang, Y. Y., Incorporation and in vitro release of doxorubicin in 

thermally sensitive micelles made from poly(N-isopropylacrylamide-co-N,N-dimethylacrylamide)b-poly(D,L-lactide-co-glycolide) 

with varying compositions, Biomaterials, 26, 5064–5074, 2005. 

123. Kabanov, A. V., Batrakova, E. V., and Miller, D. W., Pluronic block copolymers as modulators of 

drug efflux transporter activity in the blood–brain barrier, Advanced Drug Delivery Reviews, 55, 

151–164, 2003. 

124. Batrakova, E. V., Li, S., Miller, D. W., and Kabanov, A. V., Pluronic P85 increases permeability of a 

broad spectrum of drugs in polarized BBMEC and Caco-2 cell monolayers, Pharmaceutical 

Research, 16, 1366–1372, 1999. 

125. Kabanov, A. V., Batrakova, E. V., and Alakhov, V. Y., Pluronic block copolymers for overcoming 

drug resistance in cancer, Advanced Drug Delivery Reviews, 54, 759–779, 2002. 

126. Kabanov, A. V., Batrakova, E. V., Meliknubarov, N. S., Fedoseev, N. A., Dorodnich, T. Y., Alakhov, 

V. Y., Chekhonin, V. P., Nazarova, I. R., and Kabanov, V. A., A new class of drug carriers—micelles 

of poly(oxyethylene)–poly(oxypropylene) block copolymers as microcontainers for drug targeting 

from blood in brain, Journal of Controlled Release, 22, 141–157, 1992. 

127. Miller, D. W., Batrakova, E. V., and Kabanov, A. V., Inhibition of multidrug resistance-associated 

protein (MRP) functional activity with pluronic block copolymers, Pharmaceutical Research, 16, 

396–401, 1999. 

128. Rapoport, N., Marin, A. P., and Timoshin, A. A., Effect of a polymeric surfactant on electron 

transport in HL-60 cells, Archives of Biochemistry and Biophysics, 384, 100–108, 2000. 


352 


129. Slepnev, V. I., Kuznetsova, L. E., Gubin, A. N., Batrakova, E. V., Alakhov, V., and Kabanov, A. V., 

Micelles of poly(oxyethylene)–poly(oxypropylene) block copolymer (pluronic) as a tool for lowmolecular 

compound delivery into a cell: Phosphorylation of intracellular proteins with micelle 

incorporated [gamma-32P]ATP, Biochemistry International, 26, 587–595, 1992. 

130. Regev, R., Assaraf, Y. G., and Eytan, G. D., Membrane fluidization by ether, other anesthetics, and 

certain agents abolishes P-glycoprotein ATPase activity and modulates efflux from multidrug-resistant 

cells, European Journal of Biochemistry, 259, 18–24, 1999. 

131. Yamamoto, Y., Nagasaki, Y., Kato, Y., Sugiyama, Y., and Kataoka, K., Long-circulating poly(ethylene 

glycol)–poly(D,L-lactide) block copolymer micelles with modulated surface charge, Journal 


132. Poznansky, M. J. and Juliano, R. L., Biological approaches to the controlled delivery of drugs: A 

critical review, Pharmacological Reviews, 36, 277–336, 1984. 

133. Drummond, D. C., Meyer, O., Hong, K., Kirpotin, D. B., and Papahadjopoulos, D., Optimizing 

liposomes for delivery of chemotherapeutic agents to solid tumors, Pharmacological Reviews, 51, 

691–743, 1999. 

134. Patel, H. M., Serum opsonins and liposomes: Their interaction and opsonophagocytosis, Critical 

Reviews in Therapeutic Drug Carrier Systems, 9, 39–90, 1992. 

135. Moghimi, S. M. and Patel, H. M., Tissue specific opsonins for phagocytic cells and their different 

affinity for cholesterol-rich liposomes, FEBS Letters, 233, 143–147, 1988. 

136. Ulrich, F. and Zilversmit, D. B., Release from alveolar macrophages of an inhibitor of phagocytosis, 

American Journal of Physiology, 218, 1118–1127, 1970. 

137. Dunn, S. E., Brindley, A., Davis, S. S., Davies, M. C., and Illum, L., Polystyrene–poly(ethylene 

glycol) (PS–PEG2000) particles as model systems for site specific drug delivery. 2. The effect of 

PEG surface density on the in vitro cell interaction and in vivo biodistribution, Pharmaceutical 

Research, 11, 1016–1022, 1994. 

138. Batrakova, E. V., Li, S., Li, Y., Alakhov, V. Y., Elmquist, W. F., and Kabanov, A. V., Distribution 

kinetics of a micelle-forming block copolymer pluronic P85, Journal of Controlled Release, 100, 

389–397, 2004. 

139. Grindel, J. M., Jaworski, T., Piraner, O., Emanuele, R. M., and Balasubramanian, M., Distribution, 

metabolism, and excretion of a novel surface-active agent, purified poloxamer 188, in rats, dogs, and 

humans, Journal of Pharmaceutical Science, 91, 1936–1947, 2002. 

140. Novakova, K., Laznicek, M., Rypacek, F., and Machova, L., I-125-labeled PLA/PEO block copolymer: 

Biodistribution studies in rats, Journal of Bioactive and Compatible Polymers, 17, 285–296, 

2002. 

141. Ishihara, K., Nomura, H., Mihara, T., Kurita, K., Iwasaki, Y., and Nakabayashi, N., Why do phospholipid 

polymers reduce protein adsorption? Journal of Biomedical Materials Research, 39, 

323–330, 1998. 

142. Devine, D. V. and Marjan, J. M., The role of immunoproteins in the survival of liposomes in the 

circulation, Critical Reviews in Therapeutic Drug Carrier Systems, 14, 105–131, 1997. 

143. Patel, H. M. and Moghimi, S. M., Serum-mediated recognition of liposomes by phagocytic cells of 

the reticuloendothelial system—the concept of tissue specificity, Advanced Drug Delivery Reviews, 

32, 45–60, 1998. 

144. Tan, J. S., Butterfield, D. E., Voycheck, C. L., Caldwell, K. D., and Li, J. T., Surface modification of 

nanoparticles by PEO/PPO block copolymers to minimize interactions with blood components and 

prolong blood circulation in rats, Biomaterials, 14, 823–833, 1993. 


delivery, Colloids and Surfaces B: Biointerfaces, 16, 3–27, 1999. 

146. Seki, J., Sonoke, S., Saheki, A., Koike, T., Fukui, H., Doi, M., and Mayumi, T., Lipid transfer protein 

transports compounds from lipid nanoparticles to plasma lipoproteins, International Journal of 


147. Ramaswamy, M., Zhang, X., Burt, H. M., and Wasan, K. M., Human plasma distribution of free 

paclitaxel and paclitaxel associated with diblock copolymers, Journal of Pharmaceutical Science, 

86, 460–464, 1997. 



148. Mizumura, Y., Matsumura, Y., Hamaguchi, T., Nishiyama, N., Kataoka, K., Kawaguchi, T., 

Hrushesky, W. J., Moriyasu, F., and Kakizoe, T., Cisplatin-incorporated polymeric micelles eliminate 

nephrotoxicity, while maintaining antitumor activity, Japan Journal of Cancer Research, 92, 

328–336, 2001. 

149. Gianni, L., Kearns, C. M., Giani, A., Capri, G., Vigano, L., Lacatelli, A., Bonadonna, G., and Egorin, 

M. J., Nonlinear pharmacokinetics and metabolism of paclitaxel and its pharmacokinetic/pharmacodynamic 

relationships in humans, Journal of Clinical Oncology, 13, 180–190, 1995. 


drug and ultrasound, Journal of Controlled Release, 102, 203–222, 2005. 

151. Husseini, G. A., Myrup, G. D., Pitt, W. G., Christensen, D. A., and Rapoport, N. Y., Factors affecting 

acoustically triggered release of drugs from polymeric micelles, Journal of Controlled Release, 69, 

43–52, 2000. 

152. Marin, A., Muniruzzaman, M., and Rapoport, N., Mechanism of the ultrasonic activation of micellar 

drug delivery, Journal of Controlled Release, 75, 69–81, 2001. 

153. Chung, J. E., Yokoyama, M., and Okano, T., Inner core segment design for drug delivery control of 

thermo-responsive polymeric micelles, Journal of Controlled Release, 65, 93–103, 2000. 

154. Kohori, F., Sakai, K., Aoyagi, T., Yokoyama, M., Sakurai, Y., and Okano, T., Preparation and 

characterization of thermally responsive block copolymer micelles comprising poly(N-isopropylacrylamide-b-DL-lactide), 


155. Chung, J. E., Yokoyama, M., Yamato, M., Aoyagi, T., Sakurai, Y., and Okano, T., Thermo-responsive 

drug delivery from polymeric micelles constructed using block copolymers of poly(Nisopropylacrylamide) 

and poly(butylmethacrylate), Journal of Controlled Release, 62, 115–127, 

1999. 

156. Chung, J. E., Yokoyama, M., Aoyagi, T., Sakurai, Y., and Okano, T., Effect of molecular architecture 

of hydrophobically modified poly(N-isopropylacrylamide) on the formation of thermoresponsive 

core-shell micellar drug carriers, Journal of Controlled Release, 53, 119–130, 1998. 

157. Cammas, S., Suzuki, K., Sone, C., Sakurai, Y., Kataoka, K., and Okano, T., Thermo-responsive 

polymer nanoparticles with a core-shell micelle structure as site-specific drug carriers, Journal of 


158. Taillefer, J., Jones, M. C., Brasseur, N., van Lier, J. E., and Leroux, J. C., Preparation and characterization 

of pH-responsive polymeric micelles for the delivery of photosensitizing anticancer drugs, 

Journal of Pharmaceutical Science, 89, 52–62, 2000. 

159. Houlton, S., Blocking the way forward, 2003 http://www.users.globalnet.uk/~sarahx/articles/ 

mcblocks.htm. 

160. Zhang, G. D., Harada, A., Nishiyama, N., Jiang, D. L., Koyama, H., Aida, T., and Kataoka, K., 

Polyion complex micelles entrapping cationic dendrimer porphyrin: Effective photosensitizer for 

photodynamic therapy of cancer, Journal of Controlled Release, 93, 141–150, 2003. 

161. Dolmans, D. E., Fukumura, D., and Jain, R. K., Photodynamic therapy for cancer, Nature Reviews 

Cancer, 3, 380–387, 2003. 

162. van Nostrum, C. F., Polymeric micelles to deliver photosensitizers for photodynamic therapy, 


163. Jang, W. D., Nishiyama, N., Zhang, G. D., Harada, A., Jiang, D. L., Kawauchi, S., Morimoto, Y., 

Kikuchi, M., Koyama, H., Aida, T., and Kataoka, K., Supramolecular nanocarrier of anionic 

dendrimer porphyrins with cationic block copolymers modified with polyethylene glycol to 

enhance intracellular photodynamic efficacy, Angewandte Chemie International Edition in 

English, 44, 419–423, 2005. 

164. Ideta, R., Tasaka, F., Jang, W. D., Nishiyama, N., Zhang, G. D., Harada, A., Yanagi, Y., Tamaki, Y., 

Aida, T., and Kataoka, K., Nanotechnology-based photodynamic therapy for neovascular disease 

using a supramolecular nanocarrier loaded with a dendritic photosensitizer, Nano Letters, 5, 

2426–2431, 2005. 

165. Taillefer, J., Brasseur, N., van Lier, J. E., Lenaerts, V., Le Garrec, D., and Leroux, J. C., In-vitro and 

in-vivo evaluation of pH-responsive polymeric micelles in a photodynamic cancer therapy model, 

Journal of Pharmacy and Pharmacology, 53, 155–166, 2001. 


354 


166. Le Garrec, D., Taillefer, J., Van Lier, J. E., Lenaerts, V., and Leroux, J. C., Optimizing pH-responsive 

polymeric micelles for drug delivery in a cancer photodynamic therapy model, Journal of Drug 

Targeting, 10, 429–437, 2002. 

167. Woodburn, K. W., Vardaxis, N. J., Hill, J. S., Kaye, A. H., Reiss, J. A., and Phillips, D. R., Evaluation 

of porphyrin characteristics required for photodynamic therapy, Photochemistry and Photobiology, 

55, 697–704, 1992. 

168. Mamot, C., Drummond, D. C., Greiser, U., Hong, K., Kirpotin, D. B., Marks, J. D., and Park, J. W., 

Epidermal growth factor receptor (EGFR)-targeted immunoliposomes mediate specific and efficient 

drug delivery to EGFR- and EGFRvIII-overexpressing tumor cells, Cancer Research, 63, 

3154–3161, 2003. 

169. Torchilin, V. P., Lukyanov, A. N., Gao, Z. G., and Papahadjopoulos-Sternberg, B., Immunomicelles: 

Targeted pharmaceutical carriers for poorly soluble drugs, Proceedings of the National Academy of 

Sciences of the United States of America, 100, 6039–6044, 2003. 

170. Torchilin, V. P., Fluorescence microscopy to follow the targeting of liposomes and micelles to cells 

and their intracellular fate, Advanced Drug Delivery Reviews, 57, 95–109, 2005. 

171. Farokhzad, O. C., Jon, S. Y., Khadelmhosseini, A., Tran, T. N. T., LaVan, D. A., and Langer, R., 

Nanopartide-aptamer bioconjugates: A new approach for targeting prostate cancer cells, Cancer 

Research, 64, 7668–7672, 2004. 

172. Lee, H., Jang, I. H., Ryu, S. H., and Park, T. G., N-terminal site-specific mono-PEGylation of 

epidermal growth factor, Pharmaceutical Research, 20, 818–825, 2003. 

173. Mendelsohn, J. and Baselga, J., The EGF receptor family as targets for cancer therapy, Oncogene, 

19, 6550–6565, 2000. 

174. Wakebayashi, D., Nishiyama, N., Yamasaki, Y., Itaka, K., Kanayama, N., Harada, A., Nagasaki, Y., 

and Kataoka, K., Lactose-conjugated polyion complex micelles incorporating plasmid DNA as a 

targetable gene vector system: Their preparation and gene transfecting efficiency against cultured 

HepG2 cells, Journal of Controlled Release, 95, 653–664, 2004. 

175. Yasugi, K., Nakamura, T., Nagasaki, Y., Kato, M., and Kataoka, K., Sugar-installed polymer 

micelles: Synthesis and micellization of poly(ethylene glycol)–poly(D,L-lactide) block copolymers 

having sugar groups at the PEG chain end, Macromolecules, 32, 8024–8032, 1999. 

176. Jule, E., Nagasaki, Y., and Kataoka, K., Lactose-installed poly(ethylene glycol)–poly(D,L-lactide) 

block copolymer micelles exhibit fast-rate binding and high affinity toward a protein bed simulating 

a cell surface. A surface plasmon resonance study, Bioconjugate Chemistry, 14, 177–186, 2003. 

177. Lim, D. W., Yeom, Y. I., and Park, T. G., Poly(DMAEMA-NVP)-b-PEG-galactose as gene delivery 

vector for hepatocytes, Bioconjugate Chemistry, 11, 688–695, 2000. 

178. Kursa, M., Walker, G. F., Roessler, V., Ogris, M., Roedl, W., Kircheis, R., and Wagner, E., Novel 

shielded transferrin–polyethylene glycol–polyethylenimine/DNA complexes for systemic tumortargeted 

gene transfer, Bioconjugate Chemistry, 14, 222–231, 2003. 

179. Vinogradov, S., Batrakova, E., Li, S., and Kabanov, A., Polyion complex micelles with proteinmodified 

corona for receptor-mediated delivery of oligonucleotides into cells, Bioconjugate Chemistry, 

10, 851–860, 1999. 

180. Ogris, M., Walker, G., Blessing, T., Kircheis, R., Wolschek, M., and Wagner, E., Tumor-targeted 

gene therapy: Strategies for the preparation of ligand-polyethylene glycol-polyethylenimine/DNA 

complexes, Journal of Controlled Release, 91, 173–181, 2003. 

181. Lee, E. S., Na, K., and Bae, Y. H., Polymeric micelle for tumor pH and folate-mediated targeting, 


182. Merdan, T., Callahan, J., Peterson, H., Bakowsky, U., Kopeckova, P., Kissel, T., and Kopecek, J., 

Pegylated polyethylenimine-Fab 0 antibody fragment conjugates for targeted gene delivery to human 

ovarian carcinoma cells, Bioconjugate Chemistry, 14, 989–996, 2003. 

183. Kabanov, A. V. and Alakhov, V. Y., Amphiphilic Block Copolymers: Self-Assemble and Applications 

In Micelles of Amphiphilic Block Copolymers as Vehicles for Drug Delivery, Alexandris, P. 

and Lindman, B., Eds., Elsevier, The Netherlands, 1997. 

184. Lin, S. Y., Makino, K., Xia, W. Y., Matin, A., Wen, Y., Kwong, K. Y., Bourguignon, L., and Hung, 

M. C., Nuclear localization of EGF receptor and its potential new role as a transcription factor, 

Nature Cell Biology, 3, 802–808, 2001. 



185. Kircheis, R., Kichler, A., Wallner, G., Kursa, M., Ogris, M., Felzmann, T., Buchberger, M., and 

Wagner, E., Coupling of cell-binding ligands to polyethylenimine for targeted gene delivery, Gene 

Therapy, 4, 409–418, 1997. 

186. Blessing, T., Kursa, M., Holzhauser, R., Kircheis, R., and Wagner, E., Different strategies for 

formation of PEGylated EGF-conjugated PEI/DNA complexes for targeted gene delivery, Bioconjugate 

Chemistry, 12, 529–537, 2001. 

187. Gillies, E. R. and Frechet, J. M. J., Development of acid-sensitive copolymer micelles for drug 

delivery, Pure and Applied Chemistry, 76, 1295–1307, 2004. 

188. Sant, V. P., Smith, D., and Leroux, J. C., Enhancement of oral bioavailability of poorly water-soluble 

drugs by poly(ethylene glycol)-block-poly(alkyl acrylate-co-methacrylic acid) self-assemblies, 


189. Tannock, I. F. and Rotin, D., Acid pH in tumors and its potential for therapeutic exploitation, Cancer 

Research, 49, 4373–4384, 1989. 

190. Lee, E. S., Na, K., and Bae, Y. H., Super pH-sensitive multifunctional polymeric micelle, Nano 

Letters, 5, 325–329, 2005. 

191. Bae, Y., Fukushima, S., Harada, A., and Kataoka, K., Design of environment-sensitive supramolecular 

assemblies for intracellular drug delivery: Polymeric micelles that are responsive to 

intracellular pH change, Angewandte Chemie International Edition in English, 42, 4640–4643, 2003. 

192. Bae, Y., Nishiyama, N., Fukushima, S., Koyama, H., Yasuhiro, M., and Kataoka, K., Preparation and 

biological characterization of polymeric micelle drug carriers with intracellular pH-triggered drug 

release property: Tumor permeability, controlled subcellular drug distribution, and enhanced in vivo 

antitumor efficacy, Bioconjugate Chemistry, 16, 122–130, 2005. 

193. Lee, E. S., Shin, H. J., Na, K., and Bae, Y. H., Poly(L-histidine)–PEG block copolymer micelles and 

pH-induced destabilization, Journal of Controlled Release, 90, 363–374, 2003. 

194. Gao, Z. G., Lee, D. H., Kim, D. I., and Bae, Y. H., Doxorubicin loaded pH-sensitive micelle targeting 

acidic extracellular pH of human ovarian A2780 tumor in mice, Journal of Drug Targeting, 13, 

391–397, 2005. 


18 PEO-Modified 

Poly(L-Amino Acid) Micelles 

for Drug Delivery 

CONTENTS 

Xiao-Bing Xiong, Hamidreza Montazeri Aliabadi, and 

Afsaneh Lavasanifar 

18.1 Introduction ....................................................................................................................... 358 

18.2 Synthesis of PEO-b-PLAA Block Copolymers ................................................................ 358 

18.3 Micellization of PEO-b-PLAA Block Copolymers .......................................................... 361 

18.3.1 Dialysis Method .................................................................................................. 362 

18.3.2 Solvent Evaporation Method .............................................................................. 362 

18.3.3 Co-Solvent Evaporation Method ........................................................................ 362 

18.4 Rational Design and Functional Properties of PEO-b-PLAA Micelles in 

Drug Delivery .................................................................................................................... 363 

18.4.1 The PEO Shell..................................................................................................... 363 

18.4.2 The PLAA Core .................................................................................................. 363 

18.4.3 Micellar Dimensions ........................................................................................... 363 

18.4.4 Micellar Stability................................................................................................. 367 

18.4.5 Drug Incorporation and Release Properties........................................................ 368 

18.5 PEO-b-PLAA Micelles for the Delivery of Anti-Cancer Drugs ...................................... 369 

18.5.1 Cyclophosphamide .............................................................................................. 369 

18.5.2 Doxorubicin ......................................................................................................... 369 

18.5.3 Paclitaxel ............................................................................................................. 370 

18.5.4 Cisplatin and Its Derivative ................................................................................ 370 

18.5.5 Methotrexate........................................................................................................ 371 

18.5.6 KRN5500............................................................................................................. 371 

18.5.7 Camptothecin....................................................................................................... 372 

18.6 PEO-b-PLAA Micelles for the Delivery of Other Therapeutic Agents........................... 372 

18.6.1 Amphotericin B ................................................................................................... 372 

18.6.2 Indomethacin ....................................................................................................... 373 

18.7 PEO-b-PLAA Micelles for Active Drug Targeting.......................................................... 373 

18.8 PEO-b-PLAA Micelles for Gene Delivery ....................................................................... 374 

References .................................................................................................................................... 376 


357

358 


Polymeric micelles are core/shell structures formed through the self-assembly of amphiphilic block 

copolymers. The nanoscopic dimension as well as unique properties offered by separated core and 

shell domains in the structure of polymeric micelles has made them one of the most promising 

carriers for passive or active drug targeting in cancer. The nanoscopic size of polymeric micelles 

makes the carrier unrecognizable by the phagocytic cells of the reticuloendothelial system (RES), 

elongating their blood circulation, and facilitating the carrier’s extravasation from tumor vasculature. 

The small size of polymeric micelles is also expected to ease penetration of the carrier within 

the tumor tissue and further internalization of polymeric micelles into the tumor cells. Hydrophobic 

core of polymeric micelles provides an excellent host for the incorporation and stabilization of anticancer 

agents that are mostly hydrophobic. Steric effect induced by the dense hydrophilic brush on 

the micellar surface protects the carrier against attachment of proteins on the micellar surface, 

avoiding early uptake and clearance of the carrier by RES leading to prolonged circulation time and 

higher accumulation of the carrier and the encapsulated drug in selective tissues that have leaky 

vasculature (e.g., tumor or inflammation sites). Polymeric micelles have been the focus of several 

reviews in recent years. 1–11 

Among different micelle-forming block copolymers developed to date, those with poly(ethylene 

oxide) (PEO) as the shell-forming block and poly(L-amino acid)s (PLAA)s and poly(ester)s as 

the core-forming block are in the front line of drug development. The placement is owed to the 

biocompatibility of the PEO and biodegradability of PLAA and poly(ester) structures. The primary 

advantage of PEO-b-PLAA block copolymers over PEO-b-poly(esters) is the chemical flexibility of 

the PLAA structure that makes nano-engineering of the carrier a feasible approach. 4 To date, 

research on PEO-b-PLAA for drug delivery has been mainly conducted on amino acids with 

functional side groups in their chemical structure, including L-aspartic acid, L-glutamic acid, 

L-lysine, and L-histidine. The presence of free functional side groups on the PLAA block provide 

sites for the attachment of drugs, drug compatible moieties, or charged therapeutics such as DNA. 

Moreover, a systemic alteration in the structure of the core-forming block also may be used to better 

control the extent of drug loading, release, or activation. 

Chemical modifications in the structure of the PLAA block have led to the development of 

optimal PEO-b-PLAA-based polymeric micellar formulation for the delivery of a number of potent 

therapeutic agents, such as doxorubicin (DOX), 12–23 paclitaxel (PTX), 24 cisplatin (CDDP), 25–32 

amphotericin B (AmB), 33–41 etc. To this point, three PEO-b-PLAA-based polymeric micellar 

formulations (all for the solubilization and delivery of anti-cancer agents) have successfully 

passed the phase of bench-top development and advanced to the stage of clinical evaluations 

(Table 18.1). 23,24,32 This chapter provides an overview on the preparation; micellar properties 

and biological performance of PEO-modified PLAA- based polymeric micellar formulations. 

Emphasis has been placed on the application of PEO-b-PLAA- micelles for the targeted delivery 

of anti-cancer agents. At the end, new advancements in the field, including a second generation of 

PEO-b-PLAA micelles (polymeric micelles for active targeting) and development of PEO-b-PLAA 

micelles for gene delivery are briefly discussed. 

18.2 SYNTHESIS OF PEO-b-PLAA BLOCK COPOLYMERS 


The traditional method for the synthesis of PLAA relies on the ring-opening polymerization (ROP) 

of a-amino acid-N-carboxyanhydrides (NCAs) (Figure 18.1). 42 NCAs can be pre-prepared from 

a-amino acids using a solution of phosgene in THF by the Fuchs–Farthing method. 43,44 NCA 

polymerization can be initiated with different nucleophiles or bases, the most common being 

primary amines and alkoxide anions. 45–48 This synthetic method is appealing to the pharmaceutical 

industry because it involves simple reagents, and it is economical; high molecular weight polymers 

can be prepared in both good yield and large quantity by this method. 


PEO-Modified Poly(L-Amino Acid) Micelles for Drug Delivery 359 

TABLE 18.1 

Most Common PEO-Modified PLAA Micelles Investigated for Drug Delivery 

Drug Incorporation 

Method PLAA Block Incorporated Drug 

Latest Reported 

Phase of Progress References 

Chemical conjugation a 

P(Lys) Cyclophosphamid 

sulfide 

Development 96 

P(Asp) Doxorubicin Development 12, 18, 22, 98, 100 

PHEA Methotrexate Development 33, 87, 90 

Physical encapsulation b 

PBLA Doxorubicin Development 13, 94, 95 

P(Asp)-DOX Doxorubicin Clinical trials phase II 12, 17, 21, 23, 100 

PBLG Doxorubicin Development 133 

PPBA Paclitaxel Clinical trials phase I 24 

PBL-C16(Asp) KRN-5500 Development 65, 110 

PBLA Camptothecin Development 113 

P(n-butyl-L-Asp) Camptothecin Development 113 

P(lauryl-L-Asp) Camptothecin Development 113 

P(methylnaphtyl-L- 

Asp) 

Camptothecin Development 113 

PBLA Amphotericin B Development 40, 41 

PHSA Amphotericin B Development 33–39 

PBLA Indomethacin Development 64 

Electrostatic complexation c P(Asp) Cisplatin Development 27, 30, 31 

P(Glu) Cisplatin Clinical trials phase I 29, 32 

a 

Most common chemical structures shown in Figure 18.5. 

b 


c 


For the synthesis of PEO-b-PLAA block copolymers, a-methoxy-u-amino PEO has been used 

as the initiator. Synthesis of PEO-b-poly(b-benzyl-L-aspartate) (PEO-b-PBLA) and PEO-b-poly[3- 

(benzyloxycarbonyl)-L-lysine] (PEO-b-PBLL) from polymerization of b-benzyl-L-aspartate-Ncarboxyanhydride 

(BLA–NCA) and N-carboxyanhydride of 3-(benzyloxycarbonyl)-L-lysine 

(BLL–NCA), respectively, using a-methoxy-u-amino PEO as initiator has been reported. 49 

When primary amines are used as initiators, the NCA polymerization may proceed by two different 

mechanisms: amine mechanism and activated monomer (AM) mechanism (Figure 18.2A and 

Figure 18.2B). 48,50 The amine mechanism is a nucleophilic ring-opening chain-growth process 

where the polymer linearly grows with monomer conversion (Figure 18.2A). In the AM 

mechanism, NCA will be deprotonated, forming a nucleophile that initiates chain growth 

(Figure 18.2B). Polymerization by the AM mechanism may lead to side reactions. Besides, the 

initiator will not be part of the final product. In a given polymerization process, the system can 

R CH C 

NH 2 

O H N 

OH 

COCl 2 

O 

C 

O 

CH 

C 

R Nucleophile 

or base 

FIGURE 18.1 General scheme for the synthesis of PEO-b-PLAA-based block copolymers by ring-opening 

polymerization of a-amino acid-N-carboxyanhydrides (NCAs). 


O 

H 

N 

R 

O 

n

360 

(a) 

(b) 

R' NH 2 

R' NH 2 

+ 

R' NH3 N − 

R 

+ 

O 

O 

O 

O 

R 

R' 

O 

N 

H 

N 

H 

O 

R H 

N 

R 

O O 

O 

O 

R 

Reaction with NCA 

N 

or N-aminoacyl NCA 

O 

O 

O 

O R 

N-aminoacyl NCA 

n NH2 NH 2 

R 

alternate between the amine and AM mechanisms. Because a propagation step for one mechanism 

is a side reaction for the other and vice versa, block copolymers prepared from the NCA method 

using amine initiators have structures different from those predicted by monomer feed compositions 

and, most likely, have considerable homopolymer contamination. 

To avoid side reactions, transition metal initiators have been developed 51–55 that use transition 

metal complexes as the end groups to control the addition of each NCA monomer to the polymer 

chain ends. However, the presence of chain-transfer reactions prevents the preparation of high 

molecular weight PLAAs by this process. 

Living polymerization is a relatively novel method to prepare PLAA that has been developed to 

overcome the mentioned limitations of existing methods (Figure 18.3). 56–58 In this method, the 

transition metal initiator activates the monomers and forms covalent active species that permit the 

formation of polypeptides via the living polymerization of NCAs. The metals react identically with 

NCA monomers to form metallacyclic complexes by oxidative addition across the anhydride bond 

of NCA. The AB diblock, PEO-b-poly(L-lysine) (PEO-b-P(L-Lys)), ABA triblock, poly(g-benzyl- 

L-glutamate)-b-PEO-b-poly(g-benzyl-L-glutamate) (PBLG-b-PEO-b-PBLG) copolymers, and 

diblock copolymers of poly(methyl acrylate)-b-(PBLG) of high molecular weights have been 

synthesized by living polymerization mechanism. 50,56,59 

Alternatively, PEO can be coupled to PLAA after the polymerization of PLAA. For instance, 

poly(L-histidine) (P(L-His)) has been synthesized by base-initiated ROP of protected NCA of L-His 

and coupled to carboxylated PEO to form PEO-b-P(L-His) via an amide linkage using dicyclohexyl 

carbodiimide (DCC) and N-hydroxysuccinimide (NHS). 60 

R ' 

nCO 2 

H 

N 

R' NH 3 + 

N 

O 

nNCA 

O 

O 

O 

R 

R 

+ NCA 

O 

H O 

+ transfer 

−CO2 Further 

Oligopeptides 

O 

Condensation 

O 


N 

N 

H 

R' 

O 

N 

O 

O 

H 

N 

OH 

O 

R H 

N 

O O 

O 

O 

R 

H 

N 

O 

R H 

N 

R 

R 

C 

O 

n 

O 

NH 2 

O − 

R 

CO 2 

R' NH 3 + 

FIGURE 18.2 (a) Amine mechanism and (b) activated monomer (AM) mechanism of NCA polymerization. 

(From Deming, T. J., Advanced Drug Delivery Reviews, 54, 8, 2002. With permission.) 


NH 2


O 

2 NH 

(L)nM + 

3 

O 1 

4 

5 

M = Co, Ni 

R 

O 

NC A 

R 

R 

O 

−CO 2 

H 

N 

NCA 

R 

O 

M N H 

N 

H (L)n 

R 

HN 

M(L)n 

R 

n 

O 

R 

O 

N 

O 

N 

H 

O 

O 

N 

H 

R 

−nCO 2 

O-C 5 

HN 

(L) nM 

R 

(L)n M 

PEO-co-P(L-aspartic acid) (PEO-co-P(L-Asp))-bearing amine groups on the P(L-Asp) block 

have been synthesized by the melt polycondensation of N-(benzyloxycarbonyl)-L-aspartic acid 

anhydride (N-CBZ-L-ASP anhydride) and low molecular weight PEO. The product was an alternating 

copolymer having reactive amine groups on the P(L-Asp) residue. The backbone was linked 

by ester bond that is more biodegradable than the amide bond between PEO and P(L-Asp). 61 

Preparation of micelle-forming PEO-b-PLAA-based block copolymers with poly(leucine), 

poly(tyrosine), poly(phenylalanine) (PPA), and a composition peptide, i.e., poly(phenylalanineco-leucine-co-tyrosin-co-tryptophan) 

P(FLYW), as the core forming block through solid phase 

peptide synthesis has also been reported. 62 Methoxy PEO-b-PLAA dendrimers have also been 

prepared by the liquid phase peptide synthesis and used in drug delivery. 63 

18.3 MICELLIZATION OF PEO-b-PLAA BLOCK COPOLYMERS 

R 

NH 

proton migration 

O 

H 

N 

(L) nM 

FIGURE 18.3 Living polymerization of NCA. 

R 

O 

HN 

O O 

Assembly of micelle-forming PEO-b-PLAA block copolymers through one of the following 

methods has been reported. 


R 

O 

O 

O 

−CO 2 

R 

HN 

O 

O 

proton migration 

HN 

C 

M 

O 

NH 

R 

−CO 2 

R O 

N 

(L)n 

O 

N 

M 

(L)n 

R 

−CO 

H 

O 

R 

N 

R 

R 

NH 

NCA −CO 2 

O 

M 

O 

HN C 

R 

O 

(L) nM 

O 

R 

O 

N 

H 

O 

N 

H 

(L)n 

n 

NH 

N 

H 

O 

N 

O 

R 

R 

R 

O 

R 

N H 

R

362 

18.3.1 DIALYSIS METHOD 

In this method, block copolymer is dissolved in a water miscible organic solvent first. Then, this 

solution is dialyzed against water. The semi-permeable membrane keeps the micelles inside the 

dialysis bag, but it allows removal of organic solvent. Gradual replacement of the organic solvent 

with water, i.e., the non-solvent for the core-forming block, triggers self-association of block 

copolymers (Figure 18.4A). 13,35,40,64,65 

18.3.2 SOLVENT EVAPORATION METHOD 

In this process, the polymer is dissolved in a volatile organic solvent. The organic solvent is then 

removed completely by evaporation (usually under reduced pressure), leading to the formation of 

polymer films. Next, this film is reconstituted in an aqueous phase by vigorous shaking. 36,66 The 

solvent evaporation method of micellization cannot be utilized for block copolymers having large 

hydrophobic segments because the polymer film cannot be reconstituted easily in an aqueous phase 

for those structures (Figure 18.4B). 

18.3.3 CO-SOLVENT EVAPORATION METHOD 

In this approach, polymer is dissolved in a volatile water miscible organic solvent such as methanol. 

Then self-assembly is triggered by the gradual addition of aqueous phase (non-solvent for the coreforming 

block) to the organic phase. This step is followed by the evaporation of the organic 

co-solvent (Figure 18.4C). 33,34 

Polymer is dissolved 

in the organic solvent 

Gradual 

mixing with 

water 

(C) 

(A) 

(B) 

Gradual displacement of 

organic solvent with water 

through dialysis 

Solvent 

evaporation/film 

formation 


FIGURE 18.4 Micellization of PEO-b-PLAA-based block copolymers through (A) dialysis, (B) solvent 

evaporation, and (C) co-solvent evaporation methods. 



18.4 RATIONAL DESIGN AND FUNCTIONAL PROPERTIES OF PEO-b-PLAA 

MICELLES IN DRUG DELIVERY 

18.4.1 THE PEO SHELL 

PEO has a safe history of application as a pharmaceutical ingredient for the design of non-immunogenic 

peptides, carriers, or biomedical surfaces. 67,68 In PEO-b-PLAA, the presence of hydrophilic 

PEO chains introduces amphiphilicity, leading to the self-assembly of block copolymer and formation 

of association colloids. The PEO shell is expected to induce steric repulsive forces and to 

stabilize the colloidal interface against aggregation or absorption of proteins to the carrier as 

well. 10,69–71 As a result, the colloidal carrier will stay within the appropriate size range and 

surface properties that can avoid uptake by RES. A major barrier against targeted tumor delivery 

by colloidal carriers is the non-specific uptake and clearance of the carrier from the blood stream by 

RES. Avoiding RES will prolong the circulation of polymeric micelles in blood, providing more 

chance for the extravasation of the carrier at sites with leaky vasculature (e.g., solid tumors). 

The extent of steric stabilization is found to be dependent on the length and density of the PEO 

chain on the surface of colloidal carriers. 20,72–74 Elongation of the PEO chain or an increase in the 

aggregation number of PEO-b-PLAA micelles may lead to the reduced chance of micellar aggregation 

and longer circulation time in blood and a higher probability of passive targeting to the tumor 

site for polymeric micelles. 10 The same parameter may reduce the adherence and interaction of the 

polymeric micelles with target cells. 75 

18.4.2 THE PLAA CORE 

Chemical flexibility of the PLAA structure as the core-forming block is the primary advantage of 

the PEO-b-PLAA-based polymeric micelles. Existence of several functional groups on a PLAA 

block provides a number of sites for the attachment of drugs, drug compatible moieties, or charged 

therapeutics to a single polymeric backbone in micelle-forming PEO-b-PLAA block copolymers. 

This may lead to a lower dose of administration for drugs that are chemically conjugated to the 

PLAA block of PEO-b-PLAA. On the other hand, the diversity of functional groups in a PLAA 

chain (amino, hydroxyl, and carboxylic groups) allows conjugation of different chemical entities to 

the polymeric backbone and provides an opportunity for the design of polyion complex and/or pH 

responsive polymeric micelles. Finally, through changes in the chemical structure of the PLAA 

block, it will also be possible to fine tune the structure of PEO-b-PLAA micelles to achieve optimal 

properties for drug targeting. 4 From a pharmaceutical point of view, PEO-b-PLAA micelles are 

considered advantageous over PEO-b-poly(ester)s because they can easily be reconstituted after 

freeze–drying and do not need lyoprotection. 

The most effort for the production of PEO-b-PLAA micellar delivery systems has been focused 

on the application of P(L-Asp), poly(L-glutamic acid) (P(L-Glu)), and P(L-Lys)-based polymers 

(Table 18.1, Figure 18.5 through 18.7). This is owed to the presence of free carboxyl and amino 

groups that makes these structures useful for the chemical conjugation and electrostatic complexation 

of drugs and DNA to the PLAA block. 76 

A major concern for the application of PLAA-based pharmaceuticals is the possibility of immunogenic 

reactions and long-term accumulation and toxicity by these agents. Unfortunately, the 

information on the biocompatibility and biodegradability of PLAAs is limited. 77,78 Nevertheless, 

the results of ongoing clinical trials in Japan on PEO–PLAA-based micelles for the delivery of DOX, 

PTX, and CDDP is expected to provide more information on the safety of these structures. 

18.4.3 MICELLAR DIMENSIONS 

PEO-b-PLAA micelles developed to date are mostly in a diameter range of 10–100 nm. 79 At this 

specific size range, PEO-b-PLAA micelles are expected to be big enough to avoid glomerular 


364 

O 

O 

PEO-b-P(Asp)-DOX CH3 O CH2 CH2 NH 

n 

C CH 

CH2 NH 

x 

C CH2 CH NH H 

y 

C O 

C O 

R 

R = OH or 

O 

filtration and elimination by kidney and sufficiently small to escape RES clearance. 80,81 As a result, 

polymeric micelles may stay in circulation in the blood for prolonged periods, decreasing the 

clearance (CL), and volume of distribution (Vd), but increasing the area under the plasma concentration 

versus time curve (AUC) of the encapsulated drug if the drug stays within the carrier in 

biological system. This change in the pharmacokinetic profile may allow for a reduction in the dose 

of administration and/or decreased toxicity for the incorporated drug. 

Carriers of less than 200 nm with prolonged blood circulation properties are shown to passively 

accumulate in solid tumors. 82 The new blood vessels in the tumor, formed to provide access to 

nutrients and oxygen, are usually poorly made and have large gaps (150–730 nm) in their endothelium. 

This allows the extravasation of nanoparticles of the appropriate size range to the 

extravascular space surrounding the tumor cells. 83,84 Because the lymphatic system that drains 

fluids out of other organs is absent in tumors, the permeated nanocarrier is expected to become 

trapped in the tumor. This phenomenon, known as the enhanced permeation and retention (EPR) 

effect, is believed to be the reason for the passive accumulation of colloidal carriers (e.g., liposomes, 

polymerosomes, and polymeric micelles) in solid tumors. 85,86 The polymeric micelles’ 

smaller size compared to other colloidal carriers is expected to ease further penetration of the 

delivery system inside the dense environment of solid tumors. Other advantages associated with 

the small dimensions of polymeric micelles include the ease of sterilization via filtration and safety 

of administration. 

R 

OH 

OH O OH 

O 

CH3 O 

OH 

NH 

PEO-b-PHEA-MTX CH3 O CH2 CH2 CH2 m 

CH2 NH C CH NH 

CH2 C CH NH 

x 

CH2 H 

y 

HOOC CH2 CH2 CO O CH2 CH2 NH CO 

CO NH CH2 CH2 OH 

H 

NH 

C 

O 

N CH 3 

CH 2 

FIGURE 18.5 Chemical structure of some micelle-forming PEO-b-PLAA drug conjugates designed for drug 

delivery. 


O 

N 

N 

O 

N 

OH 

OH 

NH 2 


N 

O 

NH 2


O O 

PEO-b -P(L-Asp)-DOX CH3 OCH2CH2 NH 

n 

2 C CH NH C 

x 

CH2 CH2 CH NH H 

y 

C O 

C 

R 

O 

R 

O OH O 

OH 

PEO-b -PBLA 

PEO-b -P(n-butyI-L-Asp) 

PEO-b -P(lauryI-L-Asp) 

CH 3 

R = 

R = 

PEO-b -P(methynaphtyI-L-Asp) 

PEO-b -P(BLA, C16) 

R=OH or 

R = 

OH O OH 

O 

O 

CH3 OH 

NH 

O 

O CH2 CH2 NH 

n 

C CH 

CH2 NH H 

m 

CO O CH2 O 

O 

CH3 O CH2 CH2 CH2 n 

NH C CH NH C 

x 

CH2 CH NH H 

y 

CH2 C O R C O R 

CH2 CH3 3 

or H 

O 

O 

O 

O 

CH3 O CH2 CH2 CH 

n 

2 NH C CH NH 

x 

C CH2 CH NH H 

y 

CH2 CH3 11 

or H 

O 

CH C O R 

2 

O 

O 

C 

O 

O R 

CH3 O CH2 CH2 CH 

n 

2 NH C CH NH 

x 

C CH2 CH NH H 

y 

H2C or H 

CH2 CO R 

O 

C 

O 

O R 

CH3 O CH NH NH 

2 CH2 C CH C CH NH H 

n x m-x 

CH2 CH O C 

3 CH2 CH2 CH2 15 

CO O CH2 O 

FIGURE 18.6. Chemical structure of most commonly used PEO-b-PLAA-based polymers for physical encapsulation 

of drugs. 

The downside of polymeric micelles’ smaller size is the restricted space for drug encapsulation 

in the carrier. Second, this provides a limitation in the polymeric micelles’ ability for a sustained 

drug release because of a large total surface area, and finally, it allows for possible carrier accumulation 

in normal tissues bearing leaky vasculature such as the liver. Chemical engineering of the 

core structure in PEO-b-PLAA micelles may be used to modify the final dimensions of the carrier to 

an optimum scale so the polymeric micellar carrier can avoid each of these problems. Elongation of 


O 

OH 

O

366 

PEO-b -PHSA 

PEO-b -PBLG 

PEO-b -PBLG star block 

copolymer 

PEO-b -PPA 

CH3 O CH2 CH2 O CH2 S CH 

n 2 

3 

H3C CH2 C O 

NH 

CH2 C 

NH 

CH 

CH2 C 

NH 

y 

C CH NH H 

z 

CH2 C NH CH2OH 6 

16 

O 

6 

O 

O 

C 

H 3 C CH 3 


CH 3 

CH 2 

CH 2 

NH 

COO 

O 

O 

NH C CH NH 

CH2 CH2 O CH2 CH2 NH 

n 

w H 

CH 

2 2 

C O CH 

2 

C H 

6 5 

O 

CH2 

O 

O 

C CH NH 

CH 2 

CH 2 

C 

O 

H 

m 

O CH2 CH2 CH O CH2 CH2 O CH2 CH2 C O NH CH2 CH2 NH C CH NH H 

n k 

O 

O 

O 

C O CH 2 C 6H 5 

CH2 CH2 O CH2 CH2 CH O CH2CH2 O CH2 m 

COO 

O 

NH 

CH2 C O NH CH2 CH2 O 

NH C CH NH H 

s 

CH2 CH2 CH2 C O CH2C6H5 C O CH2 C6H5 CH2 CH2 2 

O 

NH C CH NH 

O 

H 

y 

CH3 O CH2 CH2 NH 

n 

O 

C 

CH 

NH 

Figure 18.6 (continued) 


H 

m 

O


O 

O 

PEO-b -P(L-Asp)-CDD P CH3 O CH2 CH2 n 

NH C CH 

CH2 NH 

x 

C CH2 CH NH 

y 

C O 

H 

R O C 

O 

PEO-b -P(L-Glu)-CDD P 

the PLAA block and an increase in the level of substitution or the length of the side chain have 

shown an increasing effect on the final size of the prepared polymeric micelle. 34,37,87 

18.4.4 MICELLAR STABILITY 

CH3 O CH2 CH2 NH C CH NH H 

n x 

CH2 R = Na + or 

FIGURE 18.7 PEO-b-PLAA-based polymers used for the formation of polyion complex micelles. 

Thermodynamically and kinetically stable polymeric micellar structures may retain their integrity 

in a biological environment for longer periods and, more effectively, avoid uptake by RES and 

elimination through the kidney and possibly change the normal organ distribution of the encapsulated 

drug the same way. PEO-b-PLAA block copolymers are reported to be thermodynamically 

stable and characterized by very low critical micellar concentrations (CMC) in mM range. 37,87,88 

Therefore, even after intravenous application and dilution in blood, PEO-b-PLAA micelles may 

still stay above CMC and in a micellar form. The chemical structure of the PLAA core may be 

modified toward more hydrophobicity to push the CMC to even lower values. 35,37 

Formation of hydrophobic, electrostatic, and hydrogen bonds in the core of PEO-b-PLAA 

micelles after assembly of block copolymers makes the micellar structure kinetically stable. As 

a result, polymeric micelles may not necessarily exist in equilibrium with polymeric unimers above 

CMC, and their dissociation will be very slow below CMC. The degree of kinetic stability is 

suggested to be high, especially for block copolymer structure with stiff or bulky core-forming 

blocks. Cross-linking of the core structure may enhance the kinetic stability of polymeric micelles 

as well. 89 

Currently, there has been no report on the development of experimental tools that can assess 

directly the in vivo stability of micellar structures. Instead, micellar stability is assessed indirectly 

through in vitro drug release and micellar dissociation studies under diluting condition of gel 

permeation chromatography (GPC). For instance, the stability of methotrexate (MTX) esters of 

PEO-block-poly(2-hydroxyethyl-L-aspartamide) (PEO-b-PHEA) (Figure 18.5) was investigated by 

GPC, in vitro. At high drug conjugation levels, MTX conjugate micelles were shown to be quite 


R 

H 3 N 

H 3 N 

O 

Pt 

C 

O 

O 

O 

Cl 

CH 2 

or 

H 3 N 

H 3 N 

Pt 

R

368 

stable and entirely elute as micelles during GPC. Likewise, the level of stearic acid substitution in 

micelles of PEO-block-poly(N-hexyl stearate-L-aspartamide) (PEO-b-PHSA) (Figure 18.6) was 

shown to influence micellar stability during in vitro studies. 87 A comparison between the pharmacokinetics 

and biodistribution of polymeric micellar carrier-encapsulated drug and free drug (or 

commercially available solubilized form of the drug) in animal models has provided an indirect 

measure for the evaluation of the carrier’s in vivo stability. For example, the presence of DOX 

dimers in PEO-b-P(Asp)-based formulation of DOX has been found to improve the pharmacokinetic 

and biodistribution profile of DOX for passive targeting. This effect has been attributed to the 

stabilization of the micellar formulation by DOX dimmers. 17 A similar effect is observed for CDDP 

polymeric micellar formulations when the structure of the core-forming block is made more 

hydrophobic by changing PLAA structure from P(L-Asp) to P(L-Glu) (Figure 18.7). 29 

18.4.5 DRUG INCORPORATION AND RELEASE PROPERTIES 


Incorporation of drugs by PEO-b-PLAA micelles has been accomplished through three different 

means: chemical conjugation of drug to the PLAA block and formation of micelle-forming polymer/drug 

conjugates; physical encapsulation of drugs in PEO-b-PLAA micelles; and formation of 

an electrostatic complex between the PLAA core and charged drug followed by the self-assembly 

of the polyion complex (Table 18.1, Figure 18.5 through 18.7). 

Drug cleavage from the polymeric backbone by non-specific hydrolysis or enzymes is a prerequisite 

for the activity of drug-block copolymer conjugates. If the micellar structure can withstand 

dissociation in a biological environment, then the hydrophobic core of polymeric micelles may 

protect the degradable bonds between the polymeric backbone and drug from exposure to water and 

hydrolysis. In this case, the micelle-forming block copolymer drug conjugate may prevent the 

premature release of the incorporated drug in blood circulation and provide a delayed mode of 

release for the cytotoxic agent that is only triggered after accumulation and uptake of the carrier by 

tumor cells. The other possible scenario will be the excessive stability of the polymeric pro-drug 

that may lead to the inactivity of the final product. Incorporation of hydrophilic moieties, formation 

of pH sensitive bonds between polymer and drug, and changes in the molecular weight of the coreforming 

block may be used to modify micellar stability and drug release properties of micelleforming 

PEO-b-PLAA-based drug conjugates. 87,90,91 

Generally, physical encapsulation of drugs within polymeric micelles is a more attractive 

approach because many drug molecules do not bear reactive functional groups for chemical conjugation. 

In this system, formation of hydrophobic interactions or hydrogen bonds between the 

micelle-forming block copolymer and physically encapsulated drug provides basis for drug solubilization 

and stabilization within the micellar carrier. Unlike polymer–drug conjugates that may 

provide a delayed release for the incorporated drug, physical encapsulation of hydrophobic drugs in 

polymeric micelles usually leads to a sustained mode of drug release from the carrier. It is possible 

that the drug content of physically loaded polymeric micelles is prematurely released in the blood 

circulation, i.e., before the carrier reaches its target. In this case, the released drug and carrier will 

follow separate paths in the biological system, and the colloidal carrier will be ineffective in terms 

of drug targeting. To lower the rate of micellar dissociation, drug diffusion, and the overall rate of 

drug release from the micellar carrier, the PLAA core of PEO-b-PLAA micelles may be tailored 

either to bear more drug-compatible moieties, to become glassy under physiological condition 

(378C), or to become cross-linked. 16,33,38,39,65,89,92–94 Finally, the method of drug incorporation 

in polymeric micelles also may be modified to improve the extent of drug loading, localization, or 

physical state of the loaded drug providing other means for controlling the rate of drug release from 

polymeric micelles. 36,66,95 

Polyion complex micelles can incorporate charged therapeutics through electrostatic 

interactions between oppositely charged polymer/drug combinations. Neutralization of the 

charge on the core-forming segment of the block copolymer will trigger self-assembly of the 



polyion complex and further stabilization of the complex within the hydrophobic environment of 

the micellar core. The extent of drug release from polyion complex micelles is dependent on the 

rate of drug exchange with free ions and proteins in the physiological media. The presence of ions 

and proteins at high concentrations in the biological environment has been shown to lead to a 

relatively fast rate of drug release and micellar instability for polyion complex micelles in vivo. In 

this context, chemical manipulation of the PLAA core may be used to increase the hydrophobicity 

and rigidity of the micellar core, restricting the penetration of free ions to the micellar core in 

polyion complex micelles. 26,29 In addition, the micellar core may be cross-linked to avoid its 

dissociation. 89,92,93 This may lead to a sustained or even delayed mode of drug release from 

the carrier. 

18.5 PEO-b-PLAA MICELLES FOR THE DELIVERY OF ANTI-CANCER DRUGS 

18.5.1 CYCLOPHOSPHAMIDE 

The first attempt for the design of PEO-b-PLAA-based micellar systems for targeted delivery of 

anti-cancer drugs was made by Ringsdorf et al. in the 1980s. 96 In this study, cyclophosphamide 

(CP) sulfide and palmitic acid were attached to the P(L-Lys) block of a PEO-b-P(L-Lys). Palmitic 

acid was used as a hydrophobic moiety to induce the required amphiphilicity for micelle-formation 

to the block copolymer. CP is an alkylating agent that is transformed to active alkylating metabolites 

via hepatic and intracellular enzymes. The polymeric micellar formulation was found to be 

efficient in the stabilization of active CP metabolite, and it caused a five-fold increase in the life 

span of L1210 tumor-bearing mice even at a reduced CP-equivalent dose. 

18.5.2 DOXORUBICIN 

The PEO-b-PLAA-based micellar formulation of DOX was developed by Kataoka et al. in the late 

1980s and early 1990s and then entered clinical trials in 2001 in Japan. In early studies, micelleforming 

PEO-b-P(L-Asp) conjugates of DOX (Figure 18.5) were synthesized and tried in different 

in vitro and in vivo cancer models for their pharmacokinetics, biodistribution, toxicity, and efficacy. 

18,22,97–100 The results of these studies pointed to a higher blood and tumor levels for PEO-b- 

P(L-Asp)-DOX micelles compared to free drug, especially for micelles having longer PEO 

chain lengths. A tumor/heart concentration ratio of 12, 8.1, and 0.9 was demonstrated for PEOb-P(L-Asp)-DOX 

with PEO chains of 12000, 5000 gmol K1 , and free DOX, respectively, 24 h after 

intravenous injection. Cardiomyopathy is the most important and dose-limiting adverse effect of 

DOX. The maximum tolerable dose (MTD) of DOX was increased (almost 20-fold) by its PEO-b- 

P(L-Asp) conjugate that allowed administration of higher drug levels. However, compared to free 

drug, higher levels of conjugated DOX were required for an equal anti-tumor activity in vitro and 

in vivo. The formulation was found to contain some un-conjugated, physically entrapped DOX that 

was, in fact, responsible for the anti-tumor activity of this system. 

In further studies, PEO-b-P(L-Asp)-DOX micelles were utilized as micellar nanocontainers for 

the physical encapsulation of DOX. 12,16,17 To delay the release of DOX unimers from the formulation 

in a biological environment, DOX dimers were co-incorporated with DOX unimer within the 

micellar core. Compared to free DOX, within 24 h, DOX encapsulated in PEO-b-P(L-Asp)-DOX 

micelles showed 28.9-fold higher AUC in plasma, 3.4-fold higher tumor AUC, less toxicity, and 

superior in vivo activity in solid and hematological cancers in mice. After 24 h, the pharmacokinetics 

of encapsulated and free DOX became similar, pointing to drug release from the 

micellar carrier. 

Because lyophilized micelles that contained DOX dimers were not reconstitutable in water after 

long periods of storage, the dimer form of DOX has been removed from this formulation for clinical 

trials. 21,23 The physically encapsulated DOX in PEO-b-P(L-Asp)-DOX micelles, i.e., NK911, 


370 

showed similar side effects to free DOX in phase I clinical trails. In pharmacokinetic evaluations in 

humans, NK911 exhibited 2.5-fold increase in half-life, 2.2-fold decrease in CL, 1.6-fold decrease 

in Vd, and 2-fold increase in plasma AUC compared to free DOX. The lower degree of change in 

pharmacokinetic parameters for the NK911 compared to liposomal formulation of DOX pointed to 

the lower stability of polymeric micellar formulation. However, infusion-related symptoms 

observed after administration of liposomal DOX were not seen for NK911. This formulation has 

entered phase II clinical trials for the treatment of metastatic pancreatic cancer. 

Another PEO-b-PLAA-based micellar structure with benzyloxy group as the core-forming 

block has also been used for physical encapsulation of DOX by the same research group. 13,94,95 

PEO-b-PBLA (Figure 18.6) efficiently encapsulated DOX and sustained its rate of release (50% 

drug release within 72 h). The formulation was not as effective as PEO-b-P(L-Asp)-DOX carrier of 

DOX in animal models in reducing the toxicity of DOX. MTD of DOX was raised 2.3-fold by PEOb-PBLA 

in C-26-bearing mice (compared to a 20-fold increase in MTD observed for the PEO-b- 

P(L-Asp)-DOX carrier). 

18.5.3 PACLITAXEL 

PTX is a poorly water-soluble drug (water solubility of 0.4 mg/mL). For clinical use, PTX is 

solubilized using a significant amount of Cremophor EL in Taxol w formulation (Bristol–Myers 

Squibb). Several attempts have been made to develop safe solubilizing agents for this important 

chemotherapeutic agent. 1 However, the problems of low solubility, drug leakage, and precipitation 

upon dilution in biological system for PTX have not been completely resolved. 

The hydrophobic nature of PTX and the absence of safe carriers that can reduce the toxicity and 

enhance the efficacy of injectable PTX have driven a tremendous amount of attention to the 

application of polymeric micellar carriers for this drug. 

The only polymeric micellar formulation showing benefit in drug targeting for PTX is its PEOb-PLAA-based 

micellar formulation, namely NK105. Recently, Hamaguchi et al. reported on the 

development of a PEO-b-PLAA-based polymeric micellar formulation for PTX delivery. 24 The 

formulation consists of physically encapsulated PTX in micelles of PEO-b-poly(4-phenyl-1-butanoate)L-aspartamide 

(PEO-b-PPBA). In preclinical studies in animal models, NK105 has shown 

86-fold increase in its AUC in plasma, 86-fold decrease in CL, and 15-fold decrease in steady-state 

volume of distribution (Vdss) compared to Taxol w after intravenous injection. 24 This has resulted 

in a 25-fold increase in drug AUC in tumor and stronger anti-tumor activity in C-26 tumor-bearing 

mice models. At a PTX equivalent dose of 100 mg/kg, a single administration of NK105 resulted in 

the disappearance of tumors, and all mice remained tumor-free thereafter. This formulation is 

currently in phase I clinical trials in Japan. 

PTX polymeric micellar formulations based on PEO-b-poly(ester) block copolymers have also 

entered clinical evaluations in Canada and South Korea. 101 Although PEO-b-poly(ester) micelles 

were remarkably successful in raising the solubility of PTX, they have mostly failed in retaining 

their drug content in biological environment, thereby regarded as ineffective in terms of drug 

targeting. 102–107 

18.5.4 CISPLATIN AND ITS DERIVATIVE 


CDDP is another anti-cancer agent that has been tried for encapsulation in polymeric micelles. 

Polyion complexes of positively charged CDDP and negatively charged free carboxyl group of the 

PEO-b-P(L-Asp) block copolymer were formed and assembled to micellar structures by Kataoka et 

al. (Figure 18.7). 26–31,76 Incorporation of CDDP in polymeric micelles has led to modest changes in 

the pharmacokinetic and biodistribution properties of CDDP in Lewis Lung carcinoma mice. 

Compared to free drug, 5.2- and 4.6-fold increases in the plasma and tumor AUC (calculated 

based on total platinum content) were observed for CDDP-loaded micelles. The low degree 



of change in pharmacokinetics of CDDP was attributed to the instability of the micellar 

formulation. 

In further studies, to restrict the ion exchange rate with the micellar core and stabilize the 

polymeric micellar formulation, the P(L-Asp) core of PEO-b-PLAA micelles was replaced with the 

more hydrophobic structure of P(L-Glu). 29 This change resulted in an increase in platinum levels in 

plasma and tumor for CDDP. In anti-tumor activity studies, four out of ten C-26-bearing mice 

receiving the PEO-b-P(L-Glu) formulation of CDDP showed complete cure, whereas, with free 

drug, no mice showed complete tumor regression. Recently, this formulation (named NC-6004) has 

concluded the stage of preclinical assessments showing 65-fold increase in plasma AUC, 9-fold 

decrease in total CL, and decreased nephro and neurotoxicity in rats. NC-6004 was also found to be 

equally effective as free CDDP against MKN-45 (human gastric cancer) tumor xenografts in mice. 

This formulation has recently moved to clinical trial in Japan. 32 

PEO-b-P(L-Glu) micelles have also been tried as carriers for a less toxic but more hydrophobic 

analog of CDDP, i.e., dichloro(1,2-diaminocyclohexane)platinum(II) (DACHPt) with hopes to 

increase the stability of polyion complex micellar formulation. In pharmacokinetic studies, a 

slightly higher platinum plasma levels for this system was observed compared to CDDP formulation, 

but tumor accumulation was similar for both drugs. 108 

18.5.5 METHOTREXATE 

Kwon et al. first reported on the conjugation of MTX to a PEO-b-P(L-Asp) derivative in 1999. 90 In 

this study, MTX was attached through an ester bond to PEO-b-PHEA, obtained by aminolysis of 

PEO-b-PBLA with ethanolamine. The yield of the reaction was greater than 90%, and MTX content 

was 9–20 wt%. The PEO-b-PHEA–MTX conjugates (Figure 18.5) showed self-assembly in an 

aqueous phase using a dialysis method that resulted in micelles with an average size of 14 nm 

and a narrow size distribution. For the in vitro release, a sample in a dialysis bag was placed into 

0.1 M phosphate buffer at varied pH (2.0–9.9). Free MTX was quickly released from the dialysis 

bag within 10 h. The in vitro release of MTX from self-assembled PEO-b-PHEA–MTX conjugate 

was slowest at pH 5 and somewhat faster at pH 2.2 and 7 (less than 20% of the drug was released 

over 260 h). At pH 10, release of MTX was rapid and almost complete in 80 h. 

In 2000, the same authors continued their study on this micellar formulation by changing the 

level of MTX substitution on the polymeric chain. 87 They obtained three different MTX substitution 

levels of 3.2, 9.0, and 19.4 wt% and showed a decrease in the CMC of PEO-b- 

PHEA–MTX with an increase in MTX substitution degree. The average diameter of micelles 

also increased from 11.7 to 27.1 nm as the level of MTX substitution was raised from 3.2 to 

19.4 wt%. The results of in vitro release experiments in PBS pH 7.4 showed a slower release 

rate at higher MTX substitution levels. For PEO-b-PHEA–MTX with 3.2, 9.0, and 19.4% of 

MTX substitution, 21, 10, and 5% of drug was released after 20 days, respectively. 

18.5.6 KRN5500 

KRN5500 is a new anti-cancer drug currently in clinical trials in Japan and the USA. It is a semisynthetic, 

water-insoluble analog of spicamycin derived from Streptomyces alanosinicus. 

KRN5500 is an inactive prodrug that is metabolized to 4-N-glycylspicamycin amino nucleoside 

(SAN-Gly) by a cytosomal enzyme. Fatty acid chains present in the chemical structure of 

KRN5500 are lost in SAN-Gly. As a result, SAN-Gly does not cross cellular membranes as 

easily as KRN5500 and is found to be 1000-fold less cytotoxic than KRN5500 in vitro. 109 

Development of a polymeric micellar formulation for KRN5500 was reported by Yokoyama 

et al. 65,110 To encapsulate KRN5500 that contains an aliphatic residue in its chemical structure, the 

aromatic group of the core-forming block in PEO-b-PBLA has been replaced with cetyl esters 

(Figure 18.6). Polymeric micellar KRN5500 and free drug were found to be similar in terms of antitumor 

activity against HT-29 (human colonic cancer) and MKN-45 xenografts in a mouse model, 


372 

but polymeric micellar formulation was less toxic. The vascular damage with fibrin clot or an 

increase in the plasma levels of BUN seen after intravenous administration of free drug was not 

observed for the polymeric micellar formulation. In a bleomycin-induced lung injury rat model, 

free KRN5500 at a dose of 3 mg/kg caused extensive pulmonary pathological changes with widespread 

hemorrhage; however, after administration of polymeric micellar formulation at a similar 

dose, no pathological change was observed, and the lung resembled the control group. 

18.5.7 CAMPTOTHECIN 

Camptothecin (CPT) is a naturally occurring cytotoxic alkaloid isolated from Camptotheca accuminata. 

It exists in two forms: biologically active lactone form and inactive carboxylate form. 111,112 

The lactone form of CPT exhibits poor aqueous solubility; therefore, it is a good drug candidate for 

incorporation in polymeric micelles. 

Camptothecin has been physically encapsulated in micelles of PEO-b-P(L-Asp) having benzyl, 

n-butyl, lauryl, and methylnaphtyl attached to the core-forming block through dialysis, emulsion, 

and solvent evaporation methods. 113 The presence of aromatic structures, e.g., benzyl or methylnaphtyl 

groups, and application of a solvent evaporation method of physical encapsulation were 

shown to be more efficient, resulting in the stabilization of encapsulated CPT in the micellar 

structures. The optimized polymeric micellar formulation of CPT was shown to stabilize the 

lactone form of CPT even in the presence of a stimulated physiological condition. In the presence 

of serum, 72% of CPT structure was protected by polymeric micelles, but free drug lost 80% of its 

active form. 

18.6 PEO-b-PLAA MICELLES FOR THE DELIVERY OF OTHER THERPEUTIC 

AGENTS 

18.6.1 AMPHOTERICIN B 


AmB is the most potent anti-fungal agent used in systemic mycosis. It is an amphiphilic drug with 

very low water solubility (aqueous solubility of 1 mg/mL). For clinical use, it is solubilized with the 

aid of a low molecular weight surfactant, sodium deoxycholate, in Fungizone w . It is also available 

as three lipid-based formulations: a lipid complex (Ablect w ), a complex with cholesteryl sulfate 

(Amphotec w ), and a conventional liposome (Ambisome w ). 

Micelles of PEO-b-PHSA (Figure 18.6) having aliphatic structures in their core were prepared 

and used for the solubilization of AmB through a solvent evaporation method. 33–36,38,39 In this 

system, the level of aliphatic chain substitution was found to affect the encapsulation efficiency and 

release rate of AmB where higher levels of aliphatic side chains in the micellar core led to higher 

AmB encapsulation and a lower rate of drug release. As a result of a sustained mode of drug release, 

hemolytic activity of AmB was reduced in its polymeric micellar formulation in comparison to free 

AmB or Fungizone w . The anti-fungal efficacy of polymeric micellar formulation was found to be 

comparable to that of free AmB. In further studies, the hemolytic activity of AmB incorporated in 

PEO-b-PLAA micelles having different fatty acid chain lengths as the core-forming block (at a 

range between 2 and 18 carbons) but similar substitution levels (86–95%) were compared. The 

results pointed to an advantage for stearic acid (saturated fatty acid with an 18 carbon chain length) 

in terms of reduced AmB hemolytic activity. In vivo efficacy studies of this formulation in a 

neutropenic murine model of disseminated candidiasis indicated a similar efficacy for PEO-b- 

PHSA formulation of AmB and Fingizone w . The physical encapsulation of AmB in PEO-b- 

PBLA at an alkaline pH has also been reported. 40,41 



18.6.2 INDOMETHACIN 

Indomethacin is a non-steroidal anti-inflammatory drug (NSAID), showing a low solubility in water 

(35 mg/mL). A properly designed polymeric micellar carrier may increase the solubility of indomethacin, 

control its renal side effects, and increase accumulation of this drug in inflamed areas. 

PEO-b-PBLA micelles have been used to solubilize indomethacin, and they showed a sustained 

mode of release for the encapsulated drug. The release rate of drug from PEO-b-PBLA micelles was 

found to be dependent on the ionization state of indomethacin where maximum control on drug 

release was observed at pH values below the pK a of indomethacin (4.5). At this condition, indomethacin 

was unionized and favored the non-polar environment of PBLA core in polymeric 

micelles. 64 

18.7 PEO-b-PLAA MICELLES FOR ACTIVE DRUG TARGETING 

The clinical advantage of polymeric micelles in terms of drug targeting may be further improved 

if the carrier can take advantage of differences between cancer and normal tissue at a cellular level. 

This may be achieved through the attachment of targeting ligands that recognize cancerous tissue 

to the surface of polymeric micelles, rendering them effective in active drug targeting. PEO-b- 

PLAA micelles are excellent candidates for active drug targeting because chemical modifications 

on both core- and shell-forming blocks are possible in PEO-b-PLAA-based carriers. Targeting 

ligands can be attached at the end of the PEO chains, and the anti-cancer drugs can be physically 

or chemically entrapped in the core. To date, various ligands such as different sugars, transferrin, 

folate residues, and peptides have been attached to polymeric micelles for active targeting. 114–117 

Park et al. developed a folate-mediated (FOL-mediated) intracellular delivery system that can 

transport therapeutic proteins or other bioactive macromolecules into specific cells that over 

express FOL receptors. 118,119 The di-block copolymer conjugate, FOL–PEO-b-P(L-Lys) 

(Figure 18.8), was physically complexed with fluorescein isothiocyanate conjugated bovine 

serum albumin (FITC–BSA) in an aqueous phase by ionic interactions. Cellular uptake of 

FOL–PEO-b-P(L-Lys)/FITC–BSA complexes was greatly enhanced against a FOL receptor over 

expressing cell line (KB cells) compared to a FOL receptor deficient cell line (A549 cells). The 

presence of an excess amount of free FOL in the medium inhibited the intracellular delivery of 

FOL–PEO-b-P(L-Lys)/FITC–BSA complexes. Therefore, the enhanced cellular uptake of 

FITC–BSA by KB cells was attributed to FOL receptor-mediated endocytosis of the complexes 

having FOL moieties on the surface. The FOL–PEO-b-P(L-Lys) di-block copolymer can potentially 

be applied for intracellular delivery of a wide range of other active agents that carry 

negative charge. 

In a separate research, the novel pH-sensitive, FOL-modified polymeric mixed micelles 

composed of FOL–PEO-b-P(L-His) and PEO-b-poly(L-lactic acid) block copolymers were 

prepared. The anti-cancer drug, DOX, was physically encapsulated in the micelles. 120 The polymeric 

micellar carrier was shown to be more efficient for the delivery of DOX to tumor cells 

in vitro, demonstrating the potential for solid tumor treatment through combining targetability 

and pH sensitivity. 

Attachment of C225, i.e., the antibody against epidermal growth factor (EGF) receptors, to the 

PEO terminus of a PEO-b-P(L-Glu)-DOX has been reported (Figure 18.9). 121 The polymeric 

immuno-conjugate C225–PEO-b-P(L-Glu)-DOX selectively bound to human vulvar squamous 

carcinoma (A431) cells that over express EGF receptors. Receptor-mediated uptake of C225– 

PEO-b-P(L-Glu)-DOX occurred rapidly (within 5 min), but non-specific uptake of PEO-b- 

P(L-Glu)-DOX required 24 h. Binding of C225–PEO-b-P(L-Glu)-DOX to A431 cells was 

blocked by pretreatment with C225 antibody. The results indicate that C225–PEO-b-P(L-Glu)- 

DOX was more potent than free DOX in inhibiting the growth of A431 cells after a 6-h 

exposure period. 


374 

H 2 N 

H 2 N 

N 

H 

N 

O 

N 

N 

FOL 

H 

N 

FOL−PEO-P(L-Lys)-CBZ 

N 

H 

N 

O 

N 

N 

HBr/OHAc 

H 

N H N 

O 

DCC/NHS 

H 

N COOH 

18.8 PEO-b-PLAA MICELLES FOR GENE DELIVERY 

O 

COOH 

COOH 

O 

C 

FOL-PEG-P(L-Lys) 

+ 

Anti-cancer drugs 

ε-CBZ-P(L-Lys) 

N 

H 

FOL−NHS 

HOOC−PEO-NH 2 

Polymer-based systems have attracted great interest for gene delivery because compared to viral 

vectors, they are simple to prepare, rather stable, easy to modify, and relatively safe. For the 

purpose of gene delivery, the polymeric carriers should bear positive charges that can interact 

with the negative charges of phosphate groups on DNA, resulting in condensation of DNA by 

the carrier. 122 The polymer/DNA complexes are usually prepared in the presence of an excess 

amount of cationic polymer, leading to a net positive charge for the complex. The net positive 

charge of the complex will increase the chance of its interaction with negatively charged cell 

membranes and facilitates cellular uptake of the polymer/DNA complex via endocytosis. P(L- 

Lys) and its derivatives have widely been investigated for gene delivery because of its high 

charge density. However, the high charge density of P(L-Lys) also results in its cytotoxicity and 

its rapid clearance from circulation after intravenous injection. PEO has been attached to P(L-Lys) 

to shield the high charge density of the P(L-Lys), reducing its toxicity and elimination from the 

body. PEO-P(L-Lys) copolymers are capable of association and micelle formation. To date, two 

types of PEO-P(L-Lys) copolymers have been used as gene carriers: AB type block copolymers, i.e., 

PEO-b-P(L-Lys) and comb-shaped PEO grafted P(L-Lys), i.e., PEO-g-P(L-Lys) (Figure 18.10). 

The PEO-b-P(L-Lys) was synthesized and investigated as a gene carrier by Seymour et al. It 

formed carriers with an average diameter of around 100 nm. 123 The cytotoxicity of the PEO-b- 

P(L-Lys)/DNA complex was found to be significantly lower than P(L-Lys)/DNA complex. PEO-b- 

O n 


O 

FOL−PEO−COOH 

O 

O 

O 

N 

H CH 

C 

n 

CH2 H2C CH2 H2C NH2 FIGURE 18.8 Synthesis of FOL–PEO-b-P(L-Lys) and formation of polyion complex micelles with charged 

anti-cancer drugs or genes. 


OH


O 

S PEO C 

O 

VS-PEO-NHS 

C225-SH 

S 

SH 

+ 

O 

O 

O N 

PEO 

O 

O 

S PEO C 

O 

O 

S PEO C 

O 

C225−PEO-P(L-Glu)-DOX 

O 

H 

N 

CO CH 

CH2 NH 

n 

CH 2 

CH 2 

CH 2 

NH 2 

O 

PEO-b-P(L-lysine) 

P(L-Glu), PBS 

pH 8 

H 

N 

VS-PEO-P(L-Glu)-DOX 

H 

C C H O 

N 

CO CH NH n 

CH 2 

CH 2 

CH 2 

CH 2 

NH 

PEO 

PEO-g-P(L-lysine) 

FIGURE 18.10 Chemical structure of PEO-b-P(L-Lys) 123–127 and PEO-g-P(L-Lys) 129 used for gene delivery. 

CH 2 

CH 2 

n 

O C NH DOX 

H 

C C H O 

N 

CH2 

CH2 

O C NH DOX 

n 

O 

O 

S PEO C 

O 

H 

N CH 

O 

C NH 

DMF 

CH 2 

CH 2 

C O 

OH 

Disopropylcabodiimide 

FIGURE 18.9 Synthesis of C225–PEO-b-P(L-Glu)-DOX and micelles with chemically loaded anti-cancer 

drugs. 


DOX 

n

376 

P(L-Lys)s were further studied by Kataoka et al. for DNA delivery. 124–127 Complexation of DNA 

and PEO-b-P(L-Lys) in their study led to the formation of polyion complex micelles with an 

average diameter of 48.5 nm. The polyion complex micelles showed higher transfection efficiency 

in human hepatoma HepG2 cells at a 4:1 positive to negative charge ratio when compared to 

P(L-Lys) of the same molecular weight. Southern blotting assay showed that naked plasmid 

DNA was degraded in the blood within 5 min after intravenous injection. On the contrary, when 

plasmid DNA was administered as complex with PEO-b-P(L-Lys) micelles, the supercoiled DNA 

was detected in the blood for 30 min. Methoxy PEO-b-P(L-Lys) dendrimers have also been studied 

as gene carriers by Choi et al., forming a spherical polymer/DNA complex at a 2:1 positive to 

negative charge ratio. 63 Multiblock copolymers of PEO-b-P(L-Lys)-g-P(L-His) with ester bonds 

between PEO and P(L-Lys) have also been investigated as gene carriers. 128 

The application of PEO-g-P(L-Lys) as gene carriers was evaluated in vitro and in vivo by Kim 

et al. PEO-g-P(L-Lys) showed a slightly lower DNA condensing effect when compared to P(L-Lys), 

but it had a 5–30-fold increase in transfection efficiency in HepG2 cells (a human liver carcinoma 

cell line). 129 PEO-g-P(L-Lys) demonstrated low cytotoxicity, early gene expression, and maintenance 

of gene expression level for up to 96 h. PEO-g-P(L-Lys) was then evaluated as a carrier of the 

anti-sense glutamic acid decarboxylase (GAD) plasmid to the pancreas for the prevention of type 1 

diabetes in mice. 130 The anti-sense mRNA was expressed in the pancreas of mouse for more than 3 

days when delivered as PEO-g-P(L-Lys) complex. 

Attachment of galactose or lactose to PEO-b-P(L-Lys) nanocarriers has been accomplished to 

target the asialoglycoprotein receptor of hepatocytes. Lactose-PEO-b-P(L-Lys) showed ten times 

higher transfection efficiency in HepG2 and lower cytotoxicity when compared to P(L-Lys)/DNA 

complex. 131 

A synthetic peptide based on apolipoprotein B100 with arterial wall-binding effects was 

selected and introduced to the end of PEO-g-P(L-Lys) for targeting the polymer/DNA complex 

to the aorta endothelial cells. 132 The arterial wall binding (AWBP) conjugate of PEO-g-P(L-Lys) 

condensed plasmid DNA and formed a spherical shape complex with a size of 100 nm. In gene 

expression studies, the transfection efficiency of the AWBP–PEO-g-P(L-Lys)/DNA complex to 

bovine aorta endothelial cells and smooth muscle cells was 150–180 times higher than that of 

P(L-Ly) or PEO-g-P(L-Lys) complexes with DNA. Free AWBP decreased the transfection efficiency 

of the AWBP–PEO-g-P(L-Lys), suggesting a targeted gene delivery by the carrier to the 

aorta endothelial cells via receptor-mediated endocytosis. 

REFERENCES 


1. Aliabadi, H. M. and Lavasanifar, A., Polymeric micelles for drug delivery, Expert Opinion on Drug 

Delivery, 3(1), 139–162, 2006. 

2. Gaucher, G., Dufresne, M. H., Sant, V. P., Kang, N., Maysinger, D., and Leroux, J., Block copolymer 

micelles: Preparation, characterization and application in drug delivery, Journal of Controlled 

Release, 109(1–3), 169–188, 2005. 

3. Kwon, G. S., Polymeric micelles for delivery of poorly water-soluble compounds, Critical Reviews 

in Therapeutic Drug Carrier Systems, 20(5), 357–403, 2003. 

4. Lavasanifar, A., Samuel, J., and Kwon, G. S., Poly(ethylene oxide)-block-poly(L-amino acid) 

micelles for drug delivery, Advanced Drug Delivery Reviews, 54(2), 169–190, 2002. 

5. Allen, C., Eisenberg, A., and Maysinger, D., Copolymer drug carriers: Conjugates, micelles and 

microspheres, STP Pharma Sciences, 9(1), 139–151, 1999. 


delivery, Colloids and Surfaces B: Biointerfaces, 16(1–4), 3–27, 1999. 

7. Kataoka, K., Harada, A., and Nagasaki, Y., Block copolymer micelles for drug delivery: Design, 

characterization and biological significance, Advanced Drug Delivery Reviews, 47(1), 113–131, 

2001. 



8. Le Garrec, D., Ranger, M., and Leroux, J. C., Micelles in anticancer drug delivery, American Journal 

of Drug Delivery, 2(1), 15–42, 2004. 

9. Jones, M. C. and Leroux, J. C., Polymeric micelles—A new generation of colloidal drug carriers, 

European Journal of Pharmaceutics and Biopharmaceutics, 48(2), 101–111, 1999. 

10. Kwon, G. S. and Kataoka, K., Block copolymer micelles as long-circulating drug vehicles, Advanced 

Drug Delivery Reviews, 16(2–3), 295–309, 1995. 

11. Torchilin, V. P., Structure and design of polymeric surfactant-based drug delivery systems, Journal 

of Controlled Release, 73(2–3), 137–172, 2001. 

12. Yokoyama, M., Okano, T., Sakurai, Y., and Kataoka, K., Improved synthesis of adriamycin-conjugated 

poly(ethylene oxide)-poly(aspartic acid) block copolymer and formation of unimodal micellar 

structure with controlled amount of physically entrapped adriamycin, Journal of Controlled Release, 

32(3), 269–277, 1994. 

13. Kwon, G. S., Naito, M., Yokoyama, M., Okano, T., Sakurai, Y., and Kataoka, K., Physical entrapment 

of adriamycin in AB block copolymer micelles, Pharmaceutical Research, 12(2), 192–195, 1995. 

14. Kwon, G., Naito, M., Yokoyama, M., Okano, T., Sakurai, Y., and Kataoka, K., Block copolymer 

micelles for drug delivery: Loading and release of doxorubicin, Journal of Controlled Release, 

48(2–3), 195–201, 1997. 

15. Kataoka, K., Matsumoto, T., Yokoyama, M., Okano, T., Sakurai, Y., Fukushima, S., Okamoto, K., 

and Kwon, G. S., Doxorubicin-loaded poly(ethylene glycol)-poly(beta-benzyl-L-aspartate) copolymer 

micelles: Their pharmaceutical characteristics and biological significance, Journal of 

Controlled Release, 64(1–3), 143–153, 2000. 



micelles and their design for in vivo delivery to a solid tumor, Journal of Controlled Release, 

50(1–3), 79–92, 1998. 

17. Yokoyama, M., Okano, T., Sakurai, Y., Fukushima, S., Okamoto, K., and Kataoka, K., Selective 

delivery of adriamycin to a solid tumor using a polymeric micelle carrier system, Journal of Drug 

Targeting, 7(3), 171–186, 1999. 

18. Kwon, G. S., Yokoyama, M., Okano, T., Sakurai, Y., and Kataoka, K., Biodistribution of micelleforming 

polymer-drug conjugates, Pharmaceutical Research, 10(7), 970–974, 1993. 

19. Yokoyama, M., Miyauchi, M., Yamada, N., Okanao, T., Sakurai, Y., Kataoka, K., and Inoue, S., 

Polymer micelles as novel drug carrier: Adriamycin-conjugated poly(ethylene glycol)–poly(aspartic 

acid) block copolymer, Journal of Controlled Release, 11(1–3), 269–278, 1990. 

20. Kwon, G., Suwa, S., Yokayama, M., Okano, T., Sakurai, Y., and Kataoka, K., Enhanced tumor 

accumulation and prolonged circulation times for micelle-forming poly(ethylene oxide-aspartate) 


21. Nakanishi, T., Fukushima, S., Okamoto, K., Suzuki, M., Matsumura, Y., Yokoyama, M., Okano, T., 

Sakurai, Y., and Kataoka, K., Development of the polymer micelle carrier system for doxorubicin, 


22. Yokoyama, M., Okano, T., Sakurai, Y., Ekimoto, H., Shibazaki, C., and Kataoka, K., Toxicity and 

antitumor activity against solid tumors of micelle-forming polymeric anticancer drug and its extremely 

long circulation in blood, Cancer Research, 51(12), 3229–3236, 1991. 

23. Matsumura, Y., Hamaguchi, T., Ura, T., Muro, K., Yamada, Y., Shimada, Y., Shirao, K., Okusaka, 

T., Ueno, H., Ikeda, M., and Watanabe, N., Phase I clinical trial and pharmacokinetic evaluation of 

NK911, a micelle-encapsulated doxorubicin, British Journal of Cancer, 91(10), 1775–1781, 2004. 

24. Hamaguchi, T., Matsumura, Y., Suzuki, M., Shimizu, K., Goda, R., Nakamura, I., Nakatomi, I., 

Yokoyama, M., Kataoka, K., and Kakizoe, T., NK105, a paclitaxel-incorporating micellar nanoparticle 

formulation, can extend in vivo antitumour activity and reduce the neurotoxicity of paclitaxel, 

British Journal of Cancer, 92, 1240–1246, 2005. 

25. Bogdanov, A. A., Martin, C., Bogdanova, A. V., Brady, T. J., and Weissleder, R., An adduct of cisdiamminedichloroplatinum(II) 

and poly(ethylene glycol)poly(L-lysine)-succinate: Synthesis and 

cytotoxic properties, Bioconjugate Chemistry, 7(1), 144–149, 1996. 

26. Nishiyama, N. and Kataoka, K., Preparation and characterization of size-controlled polymeric 

micelle containing cis-dichlorodiammineplatinum(II) in the core, Journal of Controlled Release, 

74(1–3), 83–94, 2001. 


378 


27. Nishiyama, N., Kato, Y., Sugiyama, Y., and Kataoka, K., Cisplatin-loaded polymer–metal complex 

micelle with time-modulated decaying property as a novel drug delivery system, Pharmaceutical 

Research, 18(7), 1035–1041, 2001. 

28. Nishiyama, N., Koizumi, F., Okazaki, S., Matsumura, Y., Nishio, K., and Kataoka, K., Differential 

gene expression profile between PC-14 cells treated with free cisplatin and cisplatin-incorporated 

polymeric micelles, Bioconjugate Chemistry, 14(2), 449–457, 2003. 

29. Nishiyama, N., Okazaki, S., Cabral, H., Miyamoto, M., Kato, Y., Sugiyama, Y., Nishio, K., 

Matsumura, Y., and Kataoka, K., Novel cisplatin-incorporated polymeric micelles can eradicate 

solid tumors in mice, Cancer Research, 63(24), 8977–8983, 2003. 

30. Nishiyama, N., Yokoyama, M., Aoyaga, T., Okano, T., Sakurai, Y., and Kataoka, K., Preparation and 

characterization of self-assembled polymer–metal complex micelle from cis-dichlorodiammineplatinium 

(II) and poly(ethylene glycole)–poly(aspartic acid) block copolymer in an aqueous medium, 

Langmuir, 15, 377–383, 1999. 

31. Mizumura, Y., Matsumura, Y., Hamaguchi, T., Nishiyama, N., Kataoka, K., Kawaguchi, T., 

Hrushesky, W. J., Moriyasuand, F., and Kakizoe, T., Cisplatin-incorporated polymeric micelles 

eliminate nephrotoxicity, while maintaining antitumor activity, Japanese Journal of Cancer 

Research, 92(3), 328–336, 2001. 

32. Uchino, H., Matsumura, Y., Negishi, T., Koizumi, F., Hayashi, T., Honda, T., Nishiyama, N., 

Kataoka, K., Naito, S., and Kakizoe, T., Cisplatin-incorporating polymeric micelles (NC-6004) 

can reduce nephrotoxicity and neurotoxicity of cisplatin in rats, British Journal of Cancer, 93(6), 

678–687, 2005. 

33. Adams, M. L., Andes, D. R., and Kwon, G. S., Amphotericin B encapsulated in micelles based on 

poly(ethylene oxide)-block-poly(L-amino acid) derivatives exerts reduced in vitro hemolysis but 

maintains potent in vivo antifungal activity, Biomacromolecules, 4(3), 750–757, 2003. 


encapsulated by poly(ethylene oxide)-block-poly(N-hexyl-L-aspartamide)-acyl conjugate micelles: 

Effects of acyl chain length, Journal of Controlled Release, 87(1–3), 23–32, 2003. 

35. Lavasanifar, A., Samuel, J., and Kwon, G. S., Micelles of poly(ethylene oxide)-block-poly(N-alkyl 

stearate L-aspartamide): Synthetic analogues of lipoproteins for drug delivery, Journal of Biomedical 

Materials Research, 52(4), 831–835, 2000. 

36. Lavasanifar, A., Samuel, J., and Kwon, G. S., Micelles self-assembled from poly(ethylene oxide)block-poly(N-hexyl 

stearate L-aspartamide) by a solvent evaporation method: Effect on the solubilization 

and haemolytic activity of amphotericin B, Journal of Controlled Release, 77(1–2), 155–160, 

2001. 

37. Lavasanifar, A., Samuel, J., and Kwon, G. S., The effect of alkyl core structure on micellar properties 

of poly(ethylene oxide)-block-poly(L-aspartamide) derivatives, Colloids and Surfaces B: Biointerfaces, 

22(2), 115–126, 2001. 

38. Lavasanifar, A., Samuel, J., and Kwon, G. S., The effect of fatty acid substitution on the in vitro 

release of amphotericin B from micelles composed of poly(ethylene oxide)-block-poly(N-hexyl 

stearate-L-aspartamide), Journal of Controlled Release, 79(1–3), 165–172, 2002. 

39. Lavasanifar, A., Samuel, J., Sattari, S., and Kwon, G. S., Block copolymer micelles for the encapsulation 

and delivery of amphotericin B, Pharmaceutical Research, 19(4), 418–422, 2002. 

40. Yu, B. G., Okano, T., Kataoka, K., and Kwon, G. S., Polymeric micelles for drug delivery: Solubilization 

and haemolytic activity of amphotericin B, Journal of Controlled Release, 53(1–3), 

131–136, 1998. 

41. Yu, B. G., Okano, T., Kataoka, K., and Kwon, G. S., In vitro dissociation of antifungal efficacy and 

toxicity for amphotericin B-loaded poly(ethylene oxide)-block-poly(beta-benzyl-L-aspartate) 

micelles, Journal of Controlled Release, 56(1–3), 285–291, 1998. 

42. Smeenk, J. M., Lowik, D. W. P. M., and van Hest, J. C. M., Peptide-containing block copolymers: 

Synthesis and potential applications of bio-mimetic materials, Current Organic Chemistry, 9(12), 

1115–1125, 2005. 

43. Fuller, W. D., Verlander, M. S., and Goodman, M., A procedure for the facile synthesis of aminoacid 

N-carboxyanhydrides, Biopolymers, 15(9), 1869–1871, 1976. 



44. Li, L. Y., Sun, P. C., Yao, Y., Chen, T. H., Li, B. H., Jin, Q. H., and Ding, D. T., Synthesis of 

amphiphilic diblock copolymer poly(L-alanine)-b-poly(hydroxyethyl glutamine) and its selfassembly 

in water, Chemical Journal of Chinese Universities—Chinese, 26(8), 1548–1551, 2005. 

45. Harwood, H. J., Comments concerning the mechanism of strong base initiated NCA polymerization. 

Abstract in Abstracts of papers of the American Chemical Society, 187(APR),72-Poly, 1984. 

46. Kricheldorf, H. R. and Mulhaupt, R., Mechanism of the NCA polymerization. 7. Primary and 

secondary amine-initiated polymerization of beta-amino acid NCAs, Makromolekulare Chemie- 

Macromolecular Chemistry and Physics, 180(6), 1419–1433, 1979. 

47. Kricheldorf, H. R., Von Lossow, C., and Schwarz, G., Primary amine and solvent-induced polymerizations 

of L- or D,L-phenylalanine N-carboxyanhydride, Macromolecular Chemistry and Physics, 

206(2), 282–290, 2005. 

48. Sekiguchi, H., Mechanism of N-carboxy-alpha-amino acid anhydride (NCA) polymerization, Pure 

and Applied Chemistry, 53(9), 1689–1714, 1981. 

49. Harada, A. and Kataoka, K., Formation of polyion complex micelles in an aqueous milieu from a pair 

of oppositely-charged block-copolymers with poly(ethylene glycol) segments, Macromolecules, 

28(15), 5294–5299, 1995. 

50. Deming, T. J., Methodologies for preparation of synthetic block copolypeptides: Materials with 

future promise in drug delivery, Advanced Drug Delivery Reviews, 54(8), 1145–1155, 2002. 

51. Deming, T. J., Transition metal-amine initiators for preparation of well-defined poly(gamma-benzyl 

L-glutamate), Journal of the American Chemical Society, 119(11), 2759–2760, 1997. 

52. Freireic, S., Gertner, D., and Zilkha, A., Polymerization of N-carboxy anhydrides by organotin 

catalysts, European Polymer Journal, 10(5), 439–443, 1974. 

53. Tsuruta, T., Matsuura, K., and Inoue, S., Copolymerization of propylene oxide with N-carboxy-DLalanine 

anhydride by organometallic systems, Die Makromolekulare Chemie, 83, 289–291, 1965. 

54. Yamashit, S. and Tani, H., Polymerization of gamma-benzyl L-glutamate N-carboxyanhydride with 

metal acetate-tri-normal-butylphosphine catalyst system, Macromolecules, 7(4), 406–409, 1974. 

55. Yamashit, S., Waki, K., Yamawaki, N., and Tani, H., Stereoselective polymerization of alphaamino-acid 

N-carboxyanhydrides with nickel DL-2-methylbutyrate-tri-normal-butylphosphine catalyst 

system, Macromolecules, 7(4), 410–415, 1974. 

56. Brzezinska, K. R. and Deming, T. J., Synthesis of AB diblock copolymers by atom-transfer radical 

polymerization (ATRP) and living polymerization of alpha-amino acid-N-carboxyanhydrides, 

Macromolecular Bioscience, 4(6), 566–569, 2004. 

57. Dvorak, M. and Rypacek, F., Preparation and polymerization of N-carboxyanhydrides of alphaamino-acids, 

Chemicke Listy, 89(7), 423–436, 1995. 

58. Vayaboury, W., Giani, O., Cottet, H., Deratani, A., and Schue, F., Living polymerization of alphaamino 

acid N-carboxyanhydrides (NCA) upon decreasing the reaction temperature, Macromolecular 

Rapid Communications, 25(13), 1221–1224, 2004. 

59. Deming, T. J., Living polymerization of alpha-amino acid-N-carboxyanhydrides, Journal of Polymer 

Science Part A—Polymer Chemistry, 38(17), 3011–3018, 2000. 

60. Lee, E. S., Shin, H. J., Na, K., and Bae, Y. H., Poly(L-histidine)–PEG block copolymer micelles and 

pH-induced destabilization, Journal of Controlled Release, 90(3), 363–374, 2003. 

61. Won, C. Y., Chu, C. C., and Lee, J. D., Synthesis and characterization of biodegradable poly(Laspartic 

acid-co-PEG), Journal of Polymer Science Part A—Polymer Chemistry, 36(16), 2949–2959, 

1998. 

62. Van Domeselaar, G. H., Kwon, G. S., Andrew, L. C., and Wishart, D. S., Application of solid phase 

peptide synthesis to engineering PEO-peptide block copolymers for drug delivery, Colloids and 

Surfaces B: Biointerfaces, 30(4), 323–334, 2003. 

63. Choi, J. S., Lee, E. J., Choi, Y. H., Jeong, Y. J., and Park, J. S., Poly(ethylene glycol)-block-poly(Llysine) 

dendrimer: Novel linear polymer/dendrimer block copolymer forming a spherical watersoluble 

polyionic complex with DNA, Bioconjugate Chemistry, 10(1), 62–65, 1999. 

64. La, S. B., Okano, T., and Kataoka, K., Preparation and characterization of the micelle-forming 

polymeric drug indomethacin-incorporated poly(ethylene oxide)–poly(beta-benzyl L-aspartate) 

block copolymer micelles, Journal of Pharmaceutical Sciences, 85(1), 85–90, 1996. 


380 


65. Yokoyama, M., Satoh, A., Sakurai, Y., Okano, T., Matsumura, Y., Kakizoe, T., and Kataoka, K., 

Incorporation of water-insoluble anticancer drug into polymeric micelles and control of their particle 

size, Journal of Controlled Release, 55(2–3), 219–229, 1998. 

66. Opanasopit, P., Yokoyama, M., Watanabe, M., Kawano, K., Maitani, Y., and Okano, T., Influence of 

serum and albumins from different species on stability of camptothecin-loaded micelles, Journal of 

Controlled Release, 104(2), 313–321, 2005. 

67. Working, P., Newman, M., Johnson, J., and Cornacoff, J., Safety of poly(ethylene glycol) and 

poly(ethylene glycol) derivatives, In Poly(ethylene glycol) Chemistry and Biological Applications, 

Harris, J. and Zalipsky, S., Eds., American Chemical Society, Washington, DC, pp. 44–57, 1997. 

68. Zalipsky, S. and Harris, J., Introduction to chemistry and biological applications of poly(ethylene 

glycol), In Poly(ethylene glycol) Chemistry and Biological Applications, Harris, J. and Zalipsky, S., 

Eds., American Chemical Society, Washington, DC, pp. 1–13, 1997. 

69. Stolnik, S., Illum, L., and Davis, S. S., Long circulating microparticulate drug carriers, Advanced 

Drug Delivery Reviews, 16, 195–214, 1995. 

70. Zhang, F., Kang, E., Neoh, K., and Huang, W., Modification of gold surface by grafting of poly(ethylene 

glycol) for reduction in protein adsorption and platelet adhesion, Journal of Biomaterial 

Science Polymer Edition, 12, 515–531, 2001. 

71. Trubetskoy, V. and Torchilin, V., Use of polyoxyethylene-lipid conjugates as long-circulating 

carriers for delivery of therapeutic and diagnostic agents, Advanced Drug Delivery Reviews, 16, 

311–320, 1995. 

72. Dunn, S., Brindley, A., Davis, S., Davies, M., and Illum, L., Polystyrene-poly(ethylene glycol) (PS- 

PEG) particles as model systems for site specific drug delivery. 2. The effect of PEG surface density 

on the in vitro cell interaction and in vivo biodistribution, Pharmaceutical Research, 11, 1016–1022, 

1994. 

73. Mosqueira, V. C., Legrand, P., Morgat, J. L., Vert, M., Mysiakine, E., Gref, R., Devissaguet, J. P., 

and Barratt, G., Biodistribution of long-circulating PEG-grafted nanocapsules in mice: Effects of 

PEG chain length and density, Pharmaceutical Research, 18(10), 1411–1419, 2001. 

74. Gref, R., Luck, M., Quellec, P., Marchand, M., Dellacherie, E., Harnisch, S., Blunk, T., and Muller, 

R. H., ‘Stealth’ corona-core nanoparticles surface modified by polyethylene glycol (PEG): Influences 

of the corona (PEG chain length and surface density) and of the core composition on 

phagocytic uptake and plasma protein adsorption, Colloids and Surfaces B: Biointerfaces, 

18(3–4), 301–313, 2000. 

75. Mahmud, A. and Lavasanifar, A., The effect of block copolymer structure on the internalization of 

polymeric micelles by human breast cancer cells, Colloids and Surfaces B: Biointerfaces, 45(2), 

82–89, 2005. 

76. Nishiyama, N., Bae, Y., Miyata, K., Fukushima, S., and Kataoka, K., Smart polymeric micelles for 

gene and drug delivery, Drug Discovery Today: Technologies, 2(1), 21–26, 2005. 

77. McCormick-Thomson, L., Sgouras, D., and Duncan, R., Poly(amino acid) copolymers as potential 

soluble drug delivery system. 2. Body distribution and preliminary biocompatibility testing in vitro 

and in vivo, Journal of Bioactive and Compatible Polymers, 4, 252–268, 1989. 

78. Chiu, H., Kopeckova, P., and Deshmane, A. S. K. J., Lysosomal degradability of poly(a-amino 

acids), Journal of Biomedical Materials Research, 34, 381–392, 1997. 

79. Kwon, G. S., Diblock copolymer nanoparticles for drug delivery, Critical Reviews in Therapeutic 

Drug Carrier Systems, 15(5), 481–512, 1998. 

80. Litzinger, D. C., Buiting, A. M., van Rooijen, N., and Huang, L., Effect of liposome size on the 

circulation time and intraorgan distribution of amphipathic poly(ethylene glycol)-containing liposomes, 

Biochimica et Biophysica Acta, 1190, 99–107, 1994. 

81. Papahadjopoulos, D. and Gabizon, A., Liposomes designed to avoid the reticuloendothelial system, 

Progress in Clinical and Biological Research, 343, 85–93, 1990. 

82. Charrois, G. J. R. and Allen, T. M., Rate of biodistribution of STEALTH liposomes to tumor and 

skin: Influence of liposome diameter and implications for toxicity and therapeutic activity, Biochimica 

et Biophysica Acta, 1609(1), 102–108, 2003. 

83. Jain, R. K., Delivery of molecular and cellular medicine to solid tumors, Advanced Drug Delivery 

Reviews, 46(1–3), 149–168, 2001. 



84. Dass, C., Tumour angiogenesis, vascular biology and enhanced drug delivery, Journal of Drug 

Targeting, 12(5), 245–255, 2004. 

85. Duncan, R., The dawning era of polymer therapeutics, Nature Reviews, 2, 347–360, 2003. 


effect in macromolecular therapeutics: A review, Journal of Controlled Release, 65(1–2), 271–284, 

2000. 

87. Li, Y. and Kwon, G. S., Methotrexate esters of poly(ethylene oxide)-block-poly(2-hydroxyethyl-Laspartamide). 

Part I: Effects of the level of methotrexate conjugation on the stability of micelles and 

on drug release, Pharmaceutical Research, 17(5), 607–611, 2000. 

88. Kwon, G., Naito, M., Yokoyama, M., Okano, T., Sakurai, Y., and Kataoka, K., Micelles based on AB 

block copolymers of poly(ethylene oxide) and poly(beta-benzyl L-aspartate), Langmuir, 9, 945–949, 

1993. 

89. Kakizawa, Y., Harada, A., and Kataoka, K., Glutathione-sensitive stabilization of block copolymer 

micelles composed of antisense DNA and thiolated poly(ethylene glycol)-block-poly(L-lysine): A 

potential carrier for systemic delivery of antisense DNA, Biomacromolecules, 2(2), 491–497, 2001. 

90. Li, Y. and Kwon, G. S., Micelle-like structures of poly(ethylene oxide)-block-poly(2-hydroxyethyl 

aspartamide)-methotrexate conjugates, Colloids and Surfaces B: Biointerfaces, 16, 217–226, 1999. 

91. Yokoyama, M., Kwon, G. S., Okano, T., Sakurai, Y., Naito, M., and Kataoka, K., Influencing factors 

on in vitro micelle stability of adriamycin–block copolymer conjugates, Journal of Controlled 

Release, 28(1–3), 59–65, 1994. 

92. Jaturanpinyo, M., Harada, A., Yuan, X., and Kataoka, K., Preparation of bionanoreactor based on 

core–shell structured polyion complex micelles entrapping trypsin in the core cross-linked with 

glutaraldehyde, Bioconjugate Chemistry, 15(2), 344–348, 2004. 

93. Miyata, K., Kakizawa, Y., Nishiyama, N., Harada, A., Yamasaki, Y., Koyama, H., and Kataoka, K., 

Block catiomer polyplexes with regulated densities of charge and disulfide cross-linking directed to 

enhance gene expression, Journal of the American Chemical Society, 126(8), 2355–2361, 2004. 

94. Kataoka, K., Matsumoto, T., Yokoyama, M., Okano, T., Sakurai, Y., Fukushima, S., Okamoto, K., 

and Kwon, G. S., Doxorubicin-loaded poly(ethylene glycol)–poly(beta-benzyl-L-aspartate) copolymer 

micelles: Their pharmaceutical characteristics and biological significance, Journal of 

Controlled Release, 64(1–3), 143–153, 2000. 

95. Kwon, G., Naito, M., Yokoyama, M., Okano, T., Sakurai, Y., and Kataoka, K., Block copolymer 

micelles for drug delivery: Loading and release of doxorubicin, Journal of Controlled Release, 48, 

195–201, 1997. 

96. Ringsdorf, H., Dorn, K., and Hoerpel, G., Polymeric antitumor agents on molecular and cellular 

level, In Bioactive Polymer Systems: An Overview, Gebelein, C. G. and Carraher, C. E., Eds., Plenum 

Press, New York, pp. 531–585, 1985. 

97. Kwon, G., Suwa, S., Yokoyama, M., Okano, T., Sakurai, Y., and Kataoka, K., Enhanced tumor 

accumulation and prolonged circulation times of micelle-forming poly(ethylene oxide-aspartate) 


98. Yokoyama, M., Inoue, S., Kataoka, K., Yui, N., and Sakurai, Y., Preparation of adriamycin-conjugated 

poly(ethylene glycol)–poly(aspartic acid) block copolymer: A new type of polymeric 

anticancer agent, Makromolekulare Chemie, 8, 431–435, 1987. 

99. Yokoyama, M., Kwon, G. S., Okano, T., Sakurai, Y., Seto, T., and Kataoka, K., Preparation of 

micelle forming polymer–drug conjugates, Bioconjugate Chemistry, 3, 295–301, 1992. 

100. Yokoyama, M., Miyauchi, M., Yamada, N., Okano, T., Sakurai, Y., and Kataoka, K., Characterization 

and anticancer activity of the micelle-forming polymeric anticancer drug adriamycinconjugated 

poly(ethylene glycol)–poly(aspartic acid) block copolymer, Cancer Research, 50, 

1693–1700, 1990. 

101. Kim, T. Y., Kim, D. W., Chung, J. Y., Shin, S. G., Kim, S. C., Heo, D. S., Kim, N. K., and Bang, 

Y. J., Phase I and pharmacokinetic study of Genexol-PM, a cremophor-free, polymeric micelleformulated 

paclitaxel, in patients with advanced malignancies, Clinical Cancer Research, 10(11), 

3708–3716, 2004. 

102. Kim, S. C., Kim, D. W., Shim, Y. H., Bang, J. S., Oh, H. S., Wan Kim, S., and Seo, M. H., In vivo 

evaluation of polymeric micellar paclitaxel formulation: Toxicity and efficacy, Journal of Controlled 

Release, 72(1–3), 191–202, 2001. 


382 


103. Le Garrec, D., Gori, S., Luo, L., Lessard, D., Smith, D. C., Yessine, M. A., Ranger, M., and Leroux, 

J. C., Poly(N-vinylpyrrolidone)-block-poly(D,L-lactide) as a new polymeric solubilizer for hydrophobic 

anticancer drugs: In vitro and in vivo evaluation, Journal of Controlled Release, 99(1), 

83–101, 2004. 

104. Leung, S. Y., Jackson, J., Miyake, H., Burt, H., and Gleave, M. E., Polymeric micellar paclitaxel 

phosphorylates Bcl-2 and induces apoptotic regression of androgen-independent LNCaP prostate 

tumors, Prostate, 44(2), 156–163, 2000. 

105. Zhang, X., Burt, H. M., Mangold, G., Dexter, D., Von Hoff, D., Mayer, L., and Hunter, W. L., Antitumor 

efficacy and biodistribution of intravenous polymeric micellar paclitaxel, Anti-Cancer Drugs, 

8(7), 696–701, 1997. 

106. Zhang, X., Burt, H. M., Von Hoff, D., Dexter, D., Mangold, G., Degen, D., Oktaba, A. M., and 

Hunter, W. L., An investigation of the antitumour activity and biodistribution of polymeric micellar 

paclitaxel, Cancer Chemotherapy and Pharmacology, 40(1), 81–86, 1997. 

107. Zhang, X., Jackson, J. K., and Burt, H. M., Development of amphiphilic diblock copolymers as 

micellar carriers of taxol, International Journal of Pharmaceutics, 132(1–2), 195–206, 1996. 

108. Cabral, H., Nishiyama, N., Okazaki, S., Koyama, H., and Kataoka, K., Preparation and biological 

properties of dichloro(1,2-diaminocyclohexane)platinum(II) (DACHPt)-loaded polymeric micelles, 


109. Sakai, T., Kawai, H., Kamishohara, M., Odagawa, A., Suzuki, A., Uchida, T., Kawasaki, T., Tsuruo, 

T., and Otake, N., Structure–antitumor activity relationship of semi-synthetic spicamycin derivatives, 

Journal of Antibiotics (Tokyo), 48(12), 1467–1480, 1995. 

110. Matsumura, Y., Yokoyama, M., Kataoka, K., Okano, T., Sakurai, Y., Kawaguchi, T., and Kakizoe, 

T., Reduction of the side effects of an antitumor agent, KRN5500, by incorporation of the drug into 

polymeric micelles, Japanese Journal of Cancer Research, 90(1), 122–128, 1999. 

111. Barreiro-Iglesias, R., Bromberg, L., Temchenko, M., Hatton, T. A., Concheiro, A., and Alvarez- 

Lorenzo, C., Solubilization and stabilization of camptothecin in micellar solutions of pluronic-gpoly(acrylic 

acid) copolymers, Journal of Controlled Release, 97(3), 537–549, 2004. 

112. Mi, Z. and Burke, T. G., Differential interactions of camptothecin lactone and carboxylate forms 

with human blood components, Biochemistry, 33(34), 10325–10336, 1994. 

113. Yokoyama, M., Opanasopit, P., Okano, T., Kawano, K., and Maitani, Y., Polymer design and 

incorporation methods for polymeric micelle carrier system containing water-insoluble anticancer 

agent camptothecin, Journal of Drug Targeting, 12(6), 373–384, 2004. 

114. Leamon, C. P., Weigl, D., and Hendren, R. W., Folate copolymer-mediated transfection of cultured 

cells, Bioconjugate Chemistry, 10(6), 947–957, 1999. 

115. Nagasaki, Y., Yasugi, K., Yamamoto, Y., Harada, A., and Kataoka, K., Sugar-installed polymeric 

micelle for a vehicle of an active targeting drug delivery system, Abstract in Abstracts of Papers of 

the American Chemical Society, publication no. 221, U434–U434, 2001. 

116. Nasongkla, N., Shuai, X., Ai, H., Weinberg, B. D., Pink, J., Boothman, D. A., and Gao, J. M., cRGDfunctionalized 

polymer micelles for targeted doxorubicin delivery, Angewandte Chemie-International 

Edition, 43(46), 6323–6327, 2004. 


corona for receptor-mediated delivery of oligonucleotides into cells, Bioconjugate Chemistry, 

10(5), 851–860, 1999. 

118. Kim, S. H., Jeong, J. H., Joe, C. O., and Park, T. G., Folate receptor mediated intracellular protein 

delivery using PLL–PEG–FOL conjugate, Journal of Controlled Release, 103(3), 625–634, 2005. 

119. Kim, S. H., Jeong, J. H., Chun, K. W., and Park, T. G., Target-specific cellular uptake of PLGA 

nanoparticles coated with poly(L-lysine)-poly(ethylene glycol)–folate conjugate, Langmuir, 21(19), 

8852–8857, 2005. 

120. Lee, E. S., Na, K., and Bae, Y. H., Polymeric micelle for tumor pH and folate-mediated targeting, 


121. Vega, J., Ke, S., Fan, Z., Wallace, S., Charsangavej, C., and Li, C., Targeting doxorubicin to 

epidermal growth factor receptors by site-specific conjugation of C225 to poly(L-glutamic acid) 

through a polyethylene glycol spacer, Pharmaceutical Research, 20(5), 826–832, 2003. 


compounds, Advanced Drug Delivery Reviews, 54, 203–222, 2002. 



123. Wolfert, M. A., Schacht, E. H., Toncheva, V., Ulbrich, K., Nazarova, O., and Seymour, L. W., 

Characterization of vectors for gene therapy formed by self-assembly of DNA with synthetic block 

co-polymers, Human Gene Theraphy, 7(17), 2123–2133, 1996. 

124. Katayose, S. and Kataoka, K., Remarkable increase in nuclease resistance of plasmid DNA through 

supramolecular assembly with poly(ethylene glycol)–poly(L-lysine) block copolymer, Journal of 

Pharmaceutical Sciences, 87(2), 160–163, 1998. 

125. Katayose, S. and Kataoka, K., Water-soluble polyion complex associates of DNA and poly(ethylene 

glycol)–poly(L-lysine) block copolymer, Bioconjugate Chemistry, 8(5), 702–707, 1997. 

126. Harada-Shiba, M., Yamauchi, K., Harada, A., Takamisawa, I., Shimokado, K., and Kataoka, K., 

Polyion complex micelles as vectors in gene therapy—pharmacokinetics and in vivo gene transfer, 

Gene Theraphy, 9(6), 407–414, 2002. 

127. Harada, A., Togawa, H., and Kataoka, K., Physicochemical properties and nuclease resistance of 

antisense-oligodeoxynucleotides entrapped in the core of polyion complex micelles composed of 

poly(ethylene glycol)–poly(L-lysine) block copolymers, European Journal of Pharmaceutical 

Sciences, 13(1), 35–42, 2001. 

128. Bikram, M., Ahn, C. H., Chae, S. Y., Lee, M. Y., Yockman, J. W., and Kim, S. W., Biodegradable 

poly(ethylene glycol)-co-poly(L-lysine)-g-histidine multiblock copolymers for nonviral gene 

delivery, Macromolecules, 37(5), 1903–1916, 2004. 

129. Choi, Y. H., Liu, F., Kim, J. S., Choi, Y. K., Park, J. S., and Kim, S. W., Polyethylene glycol-grafted 

poly-L-lysine as polymeric gene carrier, Journal of Controlled Release, 54(1), 39–48, 1998. 

130. Lee, M., Han, S. O., Ko, K. S., Koh, J. J., Park, J. S., Yoon, J. W., and Kim, S. W., Repression of 

GAD autoantigen expression in pancreas beta-cells by delivery of antisense plasmid/PEG-g-PLL 

complex, Molecular Therapy, 4(4), 339–346, 2001. 

131. Choi, Y. H., Liu, F., Park, J. S., and Kim, S. W., Lactose-poly(ethylene glycol)-grafted poly-L-lysine 

as hepatoma cell-targeted gene carrier, Bioconjugate Chemistry, 9(6), 708–718, 1998. 

132. Nah, J. W., Yu, L., Han, S. O., Ahn, C. H., and Kim, S. W., Artery wall binding peptide–poly(ethylene 

glycol)-grafted-poly(L-lysine)-based gene delivery to artery wall cells, Journal of 

Controlled Release, 78(1-3), 273–284, 2002. 

133. Jeong, Y. I., Nah, J. W., Lee, H. C., Kim, S. H., and Cho, C. S., Adriamycin release from flower-type 

polymeric micelle based on star-block copolymer composed of poly(gamma-benzyl L-glutamate) as 

the hydrophobic part and poly(ethylene oxide) as the hydrophilic part, International Journal of 

Pharmaceutics, 188(1), 49–58, 1999. 


19 

CONTENTS 

Hydrotropic Polymer Micelles 

for Cancer Therapeutics 

Sang Cheon Lee, Kang Moo Huh, Tooru Ooya, and 

Kinam Park 

19.1 Introduction ....................................................................................................................... 385 

19.2 General Review of Polymer Micelle Drug Carriers ......................................................... 387 

19.2.1 Preparative Methods of Polymer Micelles Based on Block Copolymers.......... 387 

19.2.2 Characterization and Properties of the Polymeric Micelles............................... 388 

19.2.3 Key Parameters for Morphology and Stabilization 

of the Polymeric Micelles ................................................................................... 389 

19.2.4 Drug Loading, Solubilization, and Drug Release............................................... 390 

19.3 Polymer Micelles for Cancer Chemotherapy ................................................................... 391 

19.3.1 Polymer Micelles as Carriers of Anti-Cancer Drugs ......................................... 391 

19.3.2 Polymer Micelles for Solubilization of Poorly Soluble Anti-Cancer Drugs ..... 391 

19.3.3 Targeting Systems Using Polymer Micelles ...................................................... 392 

19.3.4 Other Applications of Polymer Micelles for Cancer Therapy ........................... 393 

19.4 Hydrotropic Polymer Micelles .......................................................................................... 393 

19.4.1 Hydrotropy and Hydrotropic Agents .................................................................. 393 

19.4.2 Hydrotropic Agents in Pharmaceutics ................................................................ 394 

19.4.3 Mechanistic Studies and Structure–Property Relationship of Hydrotropic 

Solubilization....................................................................................................... 395 

19.4.4 Design Strategy of Hydrotropic Polymer Micelle Systems ............................... 395 

19.4.5 Hydrotropic Polymer Micelles as Carriers for Anti-Cancer Drugs ................... 400 

19.5 Conclusions and Future Perspectives................................................................................ 404 

References .................................................................................................................................... 404 


Recently, polymer micelles derived from amphiphilic block copolymers in an aqueous phase have 

attracted attention as a promising formulation for poorly water-soluble drugs. 1–4 Block copolymer 

micelles are nanosized particles with a typical core-shell structure. 5 The core solubilizes the 

hydrophobic drugs, while the corona allows the suspension of micelles in an aqueous medium. 

The use of block copolymer micelles as drug-carrying vehicles was proposed by Ringsdorf’s group 

in the 1980s. 6 The rationale for incorporating low-molecular-weight drugs into micelles is to 

overcome the problems of drug formulations such as toxic side effects, poor pharmacokinetics, 


385

386 


and limited solubility in water. Hydrophobic drugs can be solubilized into hydrophobic core 

structures of polymeric micelles, and solubilized at higher concentrations than their intrinsic 

water-solubility. 7–10 The chemical composition of polymeric micelles is attractive because it can 

be tailored to have desirable physicochemical properties for drug solubilization. Therefore, various 

amphiphilic block copolymers have been synthesized and investigated for micelle formulations of 

poorly soluble drugs. 11–14 In most polymeric micelles, hydrophobic drugs can be incorporated into 

the hydrophobic core of micelles by hydrophobic interaction and other additional interactions, such 

as the metal–ligand coordination bond, receptor–ligand interaction, and the electrostatic 

interaction. 15–17 It is believed that the more compatible drugs are with the cores of the micelles, 

the greater their ability to be dissolved. 

As drug carriers, polymeric micelles have been widely investigated for a diverse class of anticancer 

agents including paclitaxel, adriamycin, and methotrexate. 3,12,18,19 However, several limitations, 

for example, limited drug-solubilizing ability, poor stability in water after drug loading, and 

the lower stability with the higher loading content, are known to limit successful clinical 

applications. 20,21 

Hydrotropy has many advantages in enhancing the water solubility of poorly soluble drugs. 

Hydrotropes (or hydrotropic agents) self-associate to form noncovalent assemblies of nonpolar 

microdomains to solubilize poorly water-soluble solutes. 22 The high concentration of hydrotropes 

(greater than 1 M) is a key factor in enhancing water-solubility of poorly soluble solutes. 23,24 

Hydrotropes often exhibit a higher selectivity in solubilizing guest hydrophobic molecules than 

in surfactant micelles. Thus, identifying hydrotrope structures that effectively solubilize a specific 

drug molecule is important. Recently, Park et al. have examined a number of hydrotropes for 

solubilization of paclitaxel, a model poorly soluble anti-cancer agent. 25 Through screening the 

effective structures of the hydrotropes for paclitaxel solubilization, N,N-diethylnicotinamide 

(DENA) and N-picolylnicotinamide (PNA) were found to be the best hydrotropes. They increased 

the water-solubility of paclitaxel by 3–5 orders of magnitude over its normal solubility in pure 

water (about 0.3 mg/mL). They also examined polymers and hydrogels, based on DENA and PNA 

hydrotropes, to develop new polymeric solubilizing systems maintaining the benefits of hydrotropy. 

26 The hydrotropic property of the hydrotropes was maintained in their polymeric forms, and 

the highly localized concentration of the hydrotrope in polymers and hydrogels was found to be a 

main contributor to effective solubilization of paclitaxel. However, upon dilution in aqueous media, 

drugs solubilized in hydrotropic polymers precipitate due to the low physical stability of the 

formulations. Thus, a need exists for the hydrotropic formulations with a high stability. 

To overcome the limitations of current polymeric micelles, hydrotropic polymer micelle 

systems that have had core components designed with a high solubilizing capacity for poorly 

soluble anti-cancer agents, like paclitaxel, show enhanced long-term stability even at high drug 

loading. 27 The key in the polymer system designs is to introduce the identified structure of hydrotropes 

for a specific drug to the drug-solubilizing micellar cores. The benefit of using selfassembled 

structures is the congestion of hydrotropic moieties with high local concentration at 

the micellar inner core. To date, many polymeric micelles have shown limited solubilizing capacity 

for paclitaxel, and, in most cases, the maximum content of paclitaxel loaded in micelles was around 

20 wt%. 20,21 In addition to the limited loading capacity, the stability of the drug-loaded polymeric 

micelles is poor. Stability decreases as the drug loading increases. The poor colloidal stability of 

existing polymeric micelles is normally caused by the enhanced hydrophobicity of micelles after 

solubilization of paclitaxel, leading to aggregation of micelles. Because the hydrotropic moiety is 

relatively hydrophilic, it retains the hydrophilicity of the colloidal micelles even after loading of 

large amounts of paclitaxel. 

This chapter gives a general review on recent developments of polymer micelles for cancer 

chemotherapeutics, the basic concept of hydrotropic solubilization, and a systemic rationale of 

hydrotropic polymer micelle systems. In particular, it focuses on the design strategy and unique 

properties of hydrotropic micelle formulations for the delivery of poorly soluble anti-cancer drugs. 


Hydrotropic Polymer Micelles for Cancer Therapeutics 387 

19.2 GENERAL REVIEW OF POLYMER MICELLE DRUG CARRIERS 

Micelles are defined as colloidal dispersions, including a category of a dispersed system that 

consists of the particulate matter or the dispersed phase in a continuous phase or dispersion 

medium. 28,29 Micellization phenomena have been studied widely in the field of drug delivery. 

Amphiphilic small molecules have been utilized for preparing micelles, but amphiphilic macromolecules 

are better than the small molecules because of their lower critical micelle concentration 

(CMC) and higher stability. Polymeric micelles represent a separate class of micelles. Polymeric 

micelles consist of amphiphilic block copolymers which include multi-block copolymers, AB-type 

di-block copolymers, and ABA-type tri-block copolymers with a hydrophobic core (A or B 

segment) and a hydrophilic shell (A or B segment). Representative block copolymers are summarized 

in Table 19.1. 5,9,12 The amphiphilic block copolymer forms spherical micelles in aqueous 

environments with a hydrophilic shell and hydrophobic core parts (Figure 19.1). In general, the core 

part represents a molten-liquid globule and a swollen, hydrophilic corona. These micelles have a 

high solubilization capacity of poorly soluble drugs, and good potential to minimize drug 

degradation, loss, and harmful side effects. In this chapter, polymeric micelles for drug delivery 

applications are briefly summarized in terms of methods/mechanism of micelle formation, influencing 

factors of stabilization, and solubilization of poorly soluble drugs. 

19.2.1 PREPARATIVE METHODS OF POLYMER MICELLES BASED ON BLOCK COPOLYMERS 

There are two ways to prepare polymeric micelles: the direct dissolution method and the dialysis 

method (Figure 19.2). 1 One can select a method depending on the block copolymer’s solubility in 

an aqueous medium. As for direct dissolution, the block copolymer is simply dissolved in water or 

buffers at a concentration above its CMC. If needed, the temperature is elevated during the micelle 

formation. A good example of this method is poly(ethylene oxide)-block-poly(propylene oxide)block-poly(ethylene 

oxide) (PEO–PPO–PEO; Pluronics w ). 30,31 If block copolymers are not easily 

water soluble, the dialysis method is useful to prepare polymeric micelles. In this case, block 

copolymer is dissolved in a water-miscible organic solvent such as dimethylformamide (DMF), 

dimethylsulfoxide (DMSO), and acetonitrile. The mixed solution is then dialyzed against distilled 

water. During the dialysis, the organic solvent is removed by exchanging the outer medium 

(distilled water). The pore size of the semi-permeable membrane (usually SpectraPor w ) should 

be carefully selected, based on the expected size of polymeric micelles. The choice of organic 

solvent is important because the solvent effects the size and size-population distribution of polymeric 

micelles. La et al. reported that the use of DMSO as the organic solvent gave rise to PEO-bpoly(b-benzyl 

L-aspartate) (PEO–PBLA) micelles, of which size was only 17 nm. 32 However, only 


TABLE 19.1 

Representative Block Copolymers for Poorly Soluble Drugs 

Block Copolymers 

Pluronics w (PEO–PPO–PEO) 

Methoxy-PEG-b-poly(D,L-lactide) (PEG–PLA) 

PEG-b-poly(lactic acid-co-glycolic acid)-b-PEG 

Methoxy-PEG-b-polycaprolactone (PEG–PCL) 

Poly(b-maleic acid)-b-poly(b-alkylmaleic acid alkyl ester) 

Poly(N-isopropyl acrylamide)-b-PEG 

Poly(aspartic acid)-b-PEG 

PEG-b-phosphatidyl ethanolamine

388 

Hydrophilic block 

Single polymer chain 

Hydrophobic block 

Polymer micelle 

FIGURE 19.1 General scheme of micelle formation using amphiphilic block copolymers. 

6% of the PEO–PBLA formed the micelle. When dimethylacetamide (DMAc) was used as 

the organic solvent, the size was 19 nm with a high yield. Therefore, one should examine the 

effect of organic solvents on the size and size distribution, as well as the composition, of 

block copolymers. 

19.2.2 CHARACTERIZATION AND PROPERTIES OF THE POLYMERIC MICELLES 

The micellization of di- or tri-block copolymers is usually characterized by the following methods: 

1. Dye-solubilization methods using 1,6-diphenyl-1,3,5-hexatriene (DPH) and pyrene 

(Figure 19.3) 11,33–37 

2. Static and dynamic light-scattering measurements. 32,38,39 

The dye-solubilization method is useful to determine CMC, local viscosities, and polarity of the 

micellar core. 40 For example, DPH is preferentially partitioned into the hydrophobic core of 

micelles with the formation of micelles, which causes an increase in the absorbance of the dyes. 

When the concentration of the block copolymers changes, CMC can be determined as the crosspoint 

of the extrapolation of the change in absorbance over a wide concentration range. Pyrene has 

(a) 

(b) 

Dissolve block copolymer 

in aqueous medium 

(If needed, increasing temperature) 

Dissolve the copolymer 

in organic solvents 

Put into a dialysis tube 


Dialysis against water 

FIGURE 19.2 Preparation of micelles by direct dissolution (a) and dialysis methods (b). 



1,6-Diphenyl-1,3,5-Hexatriene (DPH) Pyrene 

FIGURE 19.3 Chemical structures of DPH and pyrene. 

been well-studied as a fluorescent probe to determine CMC, as well as DPH. 4,41 Below CMC, a 

small amount of pyrene is solubilized in water (the saturated concentration is as low as 6!10 K7 M). 

In the presence of polymeric micelles, pyrene is preferentially solubilized in the nonpolar micelle 

core. By increasing the concentration of the block copolymer in the presence of pyrene, an increase in 

emission intensity can be seen, at which point pyrene becomes associated with the micelle core. 

When the concentration of the block copolymers changes, CMC can be determined as the point from 

which to extrapolate the change in absorbance in a wide range of the concentration. 

Weight-average molecular weight (M W) of micelles is calculated by the results of static lightscattering 

(SLS) measurements using a Debye plot 39 : 

KðCKCMCÞ=ðRKR CMCÞ Z 1=M W C2A 2ðCKCMCÞ 

where K indicates 4p 2 ðdn=dcÞ 2 =N Al 4 , R and R CMC the excess Rayleigh ratios at concentration C and 

CMC, n is the refractive index of solution at CMC, dn/dc is the refractive index increment, N A is 

Avogadro’s number, l is the wavelength of the laser light, and A2 is the second virial coefficient. 

Hydrodynamic radii (RH) of micelles can be determined by dynamic light-scattering (DLS) 

measurements using the Stokes–Einstein equation 39,41 : 

RH Z kBT=ð3phDÞ D Z G=K 2 

where k B is the Boltzmann constant, T is the absolute temperature, h is the solvent viscosity, D is the 

diffusion coefficient obtained from the average characteristic line width (G), and K 2 is the magnitude 

of the scattering vector. Because polymeric micelles are a kind of spherical assembly, the 

diffusion coefficients should be independent of the detection angle, due to the undetectable 

rotational motions. 

19.2.3 KEY PARAMETERS FOR MORPHOLOGY AND STABILIZATION OF THE POLYMERIC MICELLES 

As for AB-type di-block copolymers, one can imagine that the di-block copolymer can form 

spherical micelles in an aqueous solution. If the hydrophilic block is too long, the di-block copolymers 

exist in water as a unimer or a macromolecular micelle. On the other hand, if the 

hydrophobic block is too long, it can show nonmicellar morphology, such as rods and lamellar 

structures. Eisenberg and his group reported that different morphologies are formed at equilibrium, 

near-equilibrium, and nonequilibrium conditions. 42 When the length of the hydrophobic segment is 

significantly shorter than that of the hydrophilic part, the block copolymer usually forms nonspherical 

micelles. The formation of “crew-cut” aggregates has been suggested by a force-balance effect 

between the hydrophobic part’s degree of stretching, the interfacial energy of the micelle core with 

water, and the interaction between hydrophilic chains as a shell. 42–44 Copolymer compositions, 


390 

concentrations, and organic solvents used for the micelle preparation significantly affect the 

morphology. 

From the pharmaceutical point of view, stability of polymeric micelles in the blood stream is 

important because administration of micelles into the body always accompanies dilution, which 

causes dissociation of the micelles. 1 The stability of polymeric micelles both in vitro and in vivo 

depends on their CMC values. However, the CMC values have been used as a thermodynamic 

stability: Equilibrium between unimers and micelles can be shifted below or above the CMC. On 

the other hand, kinetic stability, which represents the actual rate of micelle dissociation below the 

CMC, depends on the size of the hydrophobic part, the physical state of the micelle core, and 

the hydrophilic/hydrophobic ratio. The hydrophobic block plays an especially critical role. For 

example, an increase in the length of the hydrophobic block at a given length of the hydrophilic 

block causes a significant decrease in the CMC value, and an increase in the stability of the micelle. 

On the other hand, an increase in the length of hydrophilic block leads to only a small rise of CMC 

value. Additionally, CMC values of di-block copolymers are generally lower than those of tri-block 

copolymer of the same molecular weight and hydrophilic/hydrophobic ratio. Thus, the molecular 

design of block copolymers for polymeric micelles should be optimized for pharmaceutical applications. 

To increase stability, unimolecular micelles that mimic polymeric micelles regarding their 

morphological properties have been proposed since 1991. 45,46 These micelles are intrinsically 

stable upon dilution, because their formation is independent of polymer concentration. Both 

dendrimers and star polymers have been designed as unimolecular micelles. 45,47–49 

19.2.4 DRUG LOADING, SOLUBILIZATION, AND DRUG RELEASE 

Drug loading methods mostly depend on the methods of micelle formation. 50,51 In the case of 

micelles prepared by the direct dissolution in water, a stock solution of drugs in some organic 

solvents is prepared in an empty vial. Then the organic solvent is allowed to evaporate, followed by 

adding an aliquot of water containing the block copolymer. The oil-in-water emulsion method is 

another method: the drug, in a nonpolar solvent such as chloroform, is added dropwise to the water 

containing the block copolymer. In the case of the dialysis method, the drug is mixed with the block 

copolymer in common organic solvents, and the organic solvents in dialysis bags are exchanged 

with a large amount of water to induce assembly of micelles. Upon micellization, the poorly soluble 

drugs are trapped in the hydrophobic inner cores of micelles. 

It is known that the solubilization of poorly soluble drugs in the micelle core strongly depends 

on the types and efficacy of the interactions between the solubilized drug and the hydrophobic 

micelle core. 1 Drug-loading capacity depends on the compatibility between the loaded drug and the 

hydrophobic core. Drug characteristics such as polarity, hydrophobicity, and charge often affect 

compatibility; consequently, the structure and the length of hydrophobic block should be carefully 

examined to determine maximum compatibility with any poorly soluble drugs. A good parameter 

that has been used to assess compatibility is the Flory–Huggins interaction parameter, csp. This 

parameter is described by the following equation 1 : 

c sp Z ðd sKd pÞ 2 V s=RT; 


where csp is the interaction parameter between the solubilized drug (s) and the hydrophobic block 

part (p), d is the Scatchard–Hildebrand solubility parameter of the hydrophobic block part, and Vs is 

the molar volume of the solubilized drug. If the csp shows low value, the compatibility between 

drug and hydrophobic block part is great. By using this parameter, Gadelle et al. suggested the 

following mechanism for solubilization of aromatic solutes in PEO-b-PPO-b-PEO 28 : 

1. The addition of apolar solutes (drugs) promotes aggregation of the block 

copolymer molecules. 



2. The micelle core contains some water molecules. 

3. Solubilization is initially a replacement process in which water molecules are displaced 

from the micelle core by the drug. 

Drug-release kinetics generally depend upon the rate of the drug’s diffusion from the micelles. 

There are several factors that affect the release rate: Polymer–drug interactions, localization of the 

drug within the micelle, the physical state of the micelle core, the length of the hydrophobic part of 

the block copolymer, the molecular volume of the drug, and the physical state of the drug in the 

micelle. 1 For instance, if the drug is located in the micelle core as a crystal, it may act as a 

reinforcing filter. If the magnitude of interaction between the copolymer and the surface of crystallite 

is strong, the drug crystal may cause the glass-transition temperature to increase. Thus, 

solubilization of the drug, which depends on the characteristics of both the drug and the block 

copolymer, is important in designing polymeric micelles for poorly soluble drugs, such as anticancer 

drugs. 

19.3 POLYMER MICELLES FOR CANCER CHEMOTHERAPY 

19.3.1 POLYMER MICELLES AS CARRIERS OF ANTI-CANCER DRUGS 

Significant progress in cancer therapy has been made with the development of new anti-cancer 

agents and related technologies. However, the major challenge remains in delivery technologies 

that can effectively, selectively deliver anti-cancer agents to tumor sites, and avoiding systemic 

toxicity and adverse effects to normal tissues or cells. Recently, polymer micelles have attracted 

attention as a novel drug carrier system, in particular for anti-cancer agents, due to its various, 

promising properties. These nanoscaled delivery systems have been developed to demonstrate a 

series of required properties as drug carriers, such as biocompatibility, stability both in vitro and 

in vivo, targeting, and drug loading capacity. Recent research into the development of polymer 

micelle delivery system has focused on two issues, drug solubilization and targeted delivery 

system, to enhance the efficacy of the delivered drug. 52–55 

19.3.2 POLYMER MICELLES FOR SOLUBILIZATION OF POORLY SOLUBLE ANTI-CANCER DRUGS 

Because many anti-cancer drugs and drug candidates are water-insoluble or poorly water-soluble, 

applicability of solubilizing systems to these drugs is indispensable. For instance, paclitaxel is one 

of the most effective anti-cancer agents, and currently formulated with use of surfactants, including 

Cremophor EL, due to its extremely poor water solubility. 56 Despite good efficacy of the formulation, 

its clinical use is limited by serious side effects resulting from use of Cremophor EL. 57–59 

Diverse approaches—such as chemical and physical modification, use of a cosolvent, and emulsification—have 

been explored for increasing the aqueous solubility of poorly soluble drugs. 60–66 

Although several systems have shown high solubilizing effects, they have been limited by their 

stability and toxicity. Thus, development of nontoxic and stable effective solubilizing systems for 

poorly soluble drugs is important for enhancing drug efficacy with high bioavailability. 

Over the past decade polymer micelles have been extensively investigated as a potent drug 

carrier that can effectively solubilize poorly soluble anti-cancer drugs. Because polymer micelles 

have been made with biocompatible polymers such as poly(ethylene oxide), poly(D,L-lactide), etc., 

they are much less toxic than other solubilizing agents such as cosolvents or surfactants. 20,58 

Additionally, recent advances in polymer synthesis make it possible to sophisticate the design of 

polymer composition and structure so that the resulting micelle shows high solubilizing capacity 

with prolonged stability. 

Presently a number of polymer micelle systems with different compositions and structures have 

been tried to solubilize and deliver various anti-canter drugs. Several examples of the polymer 


392 

micelles used for solubilization of poorly water-soluble anti-cancer drugs are summarized in 

Table 19.2. 2–14,19,20,42,52,54 Drug-loading tests have shown that, in most polymer micelle 

systems, not only molecular design of block copolymers, but also the drug incorporation method 

are crucial factors to increase the solubilizing capacity. 1,14 

19.3.3 TARGETING SYSTEMS USING POLYMER MICELLES 

Unique core-shell structure of polymer micelle systems with a nanoscaled size may ensure 

prolonged circulation times that are favorable for passive drug targeting. 10 Furthermore, polymer 

micelles can be selectively accumulated to tumor sites by the enhanced permeability and retention 

(EPR) effect, which makes them useful as a tumor-targeting system. An alternative, passivetargeting 

strategy is to make smart micelles using stimuli-responsive amphiphilic block copolymers 

that can sense and respond to the unique tumor environment. 54,67 Kataoka et al. have developed a 

novel intracellular pH-sensitive polymeric micelle drug carrier that controls the systemic, local, and 

subcellular distributions of pharmacologically-active drugs using amphiphilic block copolymers, 

poly(ethylene glycol)–poly(aspartate hydrazone adriamycin). 68 Because the anti-cancer drug, 

adriamycin, is conjugated to the hydrophobic segments through acid-sensitive hydrazone links, 

this polymer micelle can stably preserve drugs under physiological conditions (pH 7.4), but selectively 

release the drug by sensing the intracellular pH decrease in endosomes and lysosomes (pH 5– 

6). 

The active targeting is usually achieved by attaching specific-targeting ligand molecules, such 

as monoclonal antibodies, to the micelle surface, which provides a preferential accumulation in the 

tumor sites or cells. Such ligand and other targeting moieties are introduced in other literatures. 69 

When linked with tumor-targeting ligands, polymer micelles can be used to target tumor sites with 

TABLE 19.2 

Use of Polymer Micelle Systems for Solubilization of Poorly Soluble Anti-Cancer Agents 

Polymer Composition a 


Loaded Drug Loading Content (wt%) b 

PEG–PDLLA Paclitaxel 5–25 

Methotrexate 3.7–2.8 

PEG–PDLLACL Paclitaxel 15–25 

PEG–PGACL Paclitaxel 8–11 

PEtOz–PCL Paclitaxel 0.5–7.6 

PEG–PCL Paclitaxel 4.1–20.8 

PEG–pHPMAmDL Paclitaxel 22 

PVP-b-PDLLA Paclitaxel 5 

Docetaxel 4 

Teniposide 10 

Etoposide 20 

PEG-poly(aspartate) Camptothecin 13 

PEG–P(Asp(ADR) Doxorubicin Ellipticine 8.4–7.8 

PEG–PBTMC Doxorubicin 10 

Poly(NIPAAm-co-DMAAm)-PLGA 2–8 

a 

Calculated by: (weight of loaded drug)!100/(weight of drug-loaded micelle). 

b 

PDLLA, Poly(D,L-lactide); PDLLACL, Poly(D,L-lactide-co-caprolactone); PGACL, Poly(glycolide-co-caprolactone); 

PEtOz, poly(2-ethyl-2-oxazoline); pHPMAmDL, poly(N-(2-hydroxypropyl) methacrylamide lactate); PVP, poly(N-vinylpyrrolidone); 

P(Asp(ADR), adriamycin-conjugated poly(aspartic acid) block copolymer; PBTMC, 

poly(5-benzyloxytrimethylene carbonate); NIPAAm, N-isopropylacrylamide; DMAAm, N,N-dimethylacrylamide; PLGA, 

poly(D,L-lactide-co-glycolide). 



high affinity and specificity. Recently a multifunctional polymer micelle system has been proposed 

to endow enhanced tumor selectivity and endosome-disruption property on the carrier. 68 The 

polymer micelle is designed to expose the cell interacting ligand (biotin) only under slightly 

acidic environmental conditions of various solid tumors, and show pH-dependent dissociation, 

causing the enhanced release of the loaded drug from the carrier in early endosomal pH. 

19.3.4 OTHER APPLICATIONS OF POLYMER MICELLES FOR CANCER THERAPY 

Polymer micelles have supramolecular functionalities that are not available from either individual 

polymer chains or bulk polymer solids. Based on supramolecular architecture, polymer micelles 

have much more surface area and a separate core that can be available for chemically or physically 

introducing other active agents, including optical, radioisotopic, magnetic diagnostic, and photosensitive 

agents. 70,71 Although the delivery of chemotherapeutic agents by polymer micelle 

systems is relatively well established, the application for delivering imaging and photosensitive 

agents remains in the early stages of research. 47,70,71 The polymer design and targeting strategies 

can be also utilized in delivery of other active agents for cancer treatments. Recently, photodynamic 

therapy (PDT) has been considered as a promising method of treating cancer and other diseases. 72 

PDT involves the systemic administration of photosensitizers followed by a local application of 

light, introducing photochemical reactions that generate cytotoxical substances that produce 

necrosis of a cancer cells. 70 Polymer micelles can be a good candidate for delivery of photosensitizers 

in PDT because of their high solubilizing capacity for hydrophobic drugs and potential 

targeting property. The delivery of photosensitizers by polymer micelles may allow selective 

accumulation to solid tumor tissue and resultant higher photocytotoxicity with reduced side 

effects. 47,72 Also, hydrophobic photosensitizers that are poorly soluble or tend to aggregate in 

aqueous environments can be effectively solubilized in polymer micelle system, showing high 

photodynamic efficacy by overcoming problems from intrinsic poor solubility. 70 

19.4 HYDROTROPIC POLYMER MICELLES 

19.4.1 HYDROTROPY AND HYDROTROPIC AGENTS 

Hydrotropy is a collective molecular phenomenon describing a solubilization process whereby the 

presence of large amounts of a second solute, called hydrotropes, results in an increased aqueous 

solubility of a poorly soluble compound. 22,73 Hydrotropic agents are a diverse class of substances 

with wide industrial usage, including solubilizing agents in drug formulations, agents in the separation 

of isomeric mixtures, catalysts in heterogeneous phase chemical reactions, and fillers in 

cleaning formulations and cosmetics. Some examples of hydrotropic agents are nicotinamide 

and its derivatives, anionic benzoate, benzosulfonate, neutral phenol such as catechol and resorcinol, 

aliphatic glycolsulfate, and amino acid L-proline. 22,74 Figure 19.4 shows a variety of 

compounds that are classified as hydrotropic agents. Hydrotropic agents are characterized by a 

short, bulky, and compact moiety such as an aromatic ring, whereas surfactants have long hydrocarbon 

chains. In general hydrotropic agents have a shorter hydrophobic segment, leading to a 

higher water solubility, than that of surfactants. Other examples of hydrotropic materials 

are sodium gentisate, sodium glycinate, sodium toluate, sodium naphtoate, sodium ibuprofen, 

pheniramine, lysine, tryptophan, isoniazid, and urea. 22,23 At concentrations higher than the 

minimal hydrotrope concentration, hydrotropic agents self-associate and form noncovalent 

assemblies of lowered polarity (i.e., nonpolar microdomains) to solubilize hydrophobic solutes. 

The self-aggregation of hydrotropic agents is different from surfactant self-assemblies (i.e., 

micelles), in that hydrotropes form planar or open-layer structures instead of forming compact 

spheroid assemblies. 


394 

N 

O 

NH2 

Nicotinamide Sodium nicotinate 

Sodium benzoate 

Sodium salicylate 

CH3 

CH3 

O 

OH 

O - Na + 

SO3 - Na + 

Sodium xylenesulfonate 

H2N 

19.4.2 HYDROTROPIC AGENTS IN PHARMACEUTICS 

N 

O 

OH 

O - Na + 

O 

O - Na + 

Sodium p-aminosalicylate 

Resorcinol 

Delivery of poorly water-soluble drugs by oral route has been limited by poor absorption from the 

gastrointestinal tract (GI). 75 One of the governing factors in the transport of drugs across the 

intestinal barrier is the carrier-mediated efflux mechanism, which is referred to as multidrug 

resistance (MDR). This phenomenon is mediated by the increased expression of energy-dependent 

drug efflux proteins such as P-glycoproteins, multidrug-resistance-associated proteins, and lungresistance 

proteins. Since the excretion of the absorbed drugs by this mechanism is the saturable 

process, the solubility enhancement of poorly water-soluble drugs might overcome the barrier. At 

high concentration of drugs, the efflux protein will be saturated, thereby resulting in the increase of 

the absorbed amount of drugs and bioavailability. The hydrotrope approach is a promising new 

method with great potential for poorly soluble drugs. Using hydrotropic agents is one of the easiest 

ways of increasing water-solubility of poorly soluble drugs because it only requires mixing the 

drugs with hydrotropes in water. 24,25 Hydrotropic agents are commonly used to enhance the watersolubility 

of poorly soluble drugs, and in many instances the water-solubility of these drugs is 

increased by orders of magnitude. The use of hydrotropic agents offers numerous benefits over 

other solubilization methods such as micellar solubilization, cosolvency, and salting-in. 74 To date, 

various hydrotropic agents have been utilized to enhance the aqueous solubility of many hydrophobic 

drugs including paclitaxel, ketoprofen, diazepam, allopurinol, indomethacin, and 

griseofulvin. In most cases, the high concentration of hydrotropes is key to increasing water 

solubility of poorly soluble drugs. Thus, the approach to increase the local concentration of hydrotropes 

may provide an opportunity to develop new hydrotrope-based solubilizing systems. Despite 

the advantages of hydrotropes, application of low molecular weight hydrotropes in drug delivery 

has not been practical because it may result in absorption of a significant amount of hydrotropes 

themselves into the body, along with the drug. Recently polymeric forms of hydrotropes, such as 

OH 

CH3 

FIGURE 19.4 Chemical structures of hydrotropes used in the literature. 



O 

O - Na + 

SO3 - Na + 

Sodium p-toluenesulfonate 

H 3C (CH 2) 3 OCH 2CH 2 OSO 3 - Na + 

Sodium butyl monoglycol sulfate


hydrotropic polymers and hydrogels, were developed to overcome the drug formulations based on 

low molecular weight hydrotropes. 26 The polymeric hydrotropes were found to maintain the hydrotropic 

activity of its low molecular weight hydrotropes. Their solubilizing ability for paclitaxel was 

well examined, and the solubilization mechanism correlated with the assembly of polymeric hydrotropes. 

However, the low physical stability of drug formulations is also reported to limit the 

practical applications of these formulations. 

19.4.3 MECHANISTIC STUDIES AND STRUCTURE–PROPERTY RELATIONSHIP OF HYDROTROPIC 

SOLUBILIZATION 

The term hydrotropy does not mean a specific mechanism, but implies a collective solubilization 

phenomenon that is incompletely understood. There have been various theoretical and experimental 

studies aimed at explaining hydrotropic solubilization. Most of proposed mechanisms 

can be classified into following two schemes. The first one involves the complex formation 

between the hydrotrope and the solute. As a representative example, nicotinamide—a nontoxic 

vitamin B 3—has been shown to enhance the solubilities of a wide variety of hydrophobic drugs 

through complexation. Using nicotinamide and its derivatives, such as N-methylnicotinamide and 

N,N-diethylnicotinamide (DENA), Rasool et al. described the aromaticity of pyridine ring, which 

might promote the stacking of molecules through its planarity, as a most significant contributor in 

complexation, because the abilities of aromatic amide ligands to enhance the aqueous solubilities of 

tested drugs was higher than those of the aliphatic amide ligands. 76 The other mechanism for 

hydrotropic solubilization is self-association of the hydrotrope in an aqueous phase. This view is 

supported by experimental data, proving that some hydrotropes, including nicotinamide and 

aromatic sulfonates, associate in aqueous solutions. Using nicotinamide-riboflavin system, 

Kildsig et al. showed that the self-association of nicotinamide contributed to the solubility increase 

of riboflavin, rather than complexation between two species. 77 

Each hydrotropic agent is effective in increasing the water solubility of selected hydrophobic 

drugs, and no hydrotropic agents were found that are universally effective. Therefore, the structure– 

activity relationship between selected pairs of the hydrotropic agent and the drug is another 

important concern to discuss. Park et al. reported on the intensive studies of the elucidation of 

structure–activity relationship for the hydrotropic solubilization of paclitaxel using not only nicotinamide 

and its analogues as aromatic amides, but also various ureas as aliphatic amides. 25 They 

showed that nicotinamide enhanced the aqueous solubility of paclitaxel to a greater extent than the 

aliphatic analogues of nicotinamide such as nipecotamide and N,N-dimethylacetamide. In addition, 

the aqueous solubility of paclitaxel was found to be strongly dependent on the alkyl substituent on 

the amide nitrogen of nicotinamide. DENA of 3.5 M enhanced the aqueous solubility of paclitaxel 

up to 39.07 mg/mL, whereas N,N-dimethylnicotinamide showed paclitaxel solubility of 

1.77 mg/mL at the same concentration. They synthesized new hydrotropic agents based on the 

structure of nicotinamide by varying the substituent to the amide nitrogen with a pyridine ring or an 

ally group. The aqueous solution of N-picolylnicotinamide (PNA) (3.5 M), having another pyridine 

ring as a substituent on the amide nitrogen, was highly effective in the enhancement of paclitaxel 

solubility (29.44 mg/mL), compared with the hydrotropic effect of nicotinamide on the 

paclitaxel (0.69 mg/mL). Besides, N-allylnicotinamide highly contributed to the enhancement of 

paclitaxel solubility (14.18 mg/mL) compared to that of N,N-dimethylnicotianmide (1.77 mg/mL). 

The most interesting find is that the threshold concentration where the association occurred is 

consistent with the threshold concentration where paclitaxel solubility began to increase. 

19.4.4 DESIGN STRATEGY OF HYDROTROPIC POLYMER MICELLE SYSTEMS 

Hydrotropic polymer micelles were firstly developed based on the self-assemblies of amphiphilic 

block copolymers containing the hydrophilic block, and the hydrophobic block containing pendent 


396 

hydrotropic moieties. The main strategy for hydrotropic systems is to maximize the local concentration 

of hydrotropic structure because the drug solubilizing capacity increases when increasing 

the local concentration of hydrotropes. 26 The hydrophobic block possesses hydrotropes assembled 

in the limited volume in micellar core region. Considering several hundreds of aggregation numbers 

of micelles, the local concentration of hydrotropes would be maximized as compared to other 

polymeric hydrotropes. In addition to the enhanced solubilizing ability for poorly soluble drugs, 

the intrinsic hydrophilicity of hydrotropic structure can retain the colloidal stability by compensation 

of enhanced hydrophobicity of drug-loaded polymer micelles. Using atom-transfer radical 

polymerization (ATRP) of modified hydrotropes, Park et al. started the synthesis of amphiphilic 

block copolymers consisting of PEG and poly(2-(4-(vinylbenzyloxyl)-N,N-diethylnicotinamide) 

(PDENA) or poly(2-(4-(vinylbenzyloxyl)-N-picolylnicotinamide) (P(2-VBOPNA)) as shown in 

Figure 19.5. 27 

PEG-b-PDENA copolymers associated to form polymeric micelles with the hydrotrope-rich 

core and the PEG outer shell in aqueous media. The mean hydrodynamic diameters were in the 

range of 30–100 nm. In these polymer systems, the solubilization of paclitaxel is based on a synergistic 

effect of the unique micelle characteristics and hydrotropic activity. Hydrotropic micelles 

demonstrated not only higher loading capacity (up to 37 wt% of paclitaxel) but also enhanced 

physical stability in aqueous media. 27 The enhanced stability is due to the intrinsic hydrophilicity 

of hydrotropic moieties in the micellar core. The drug loading into the hydrotropic polymeric micelle 

core is mainly based on the attractive interactions between the hydrotropic moiety and paclitaxel. 

The structure of hydrotropic polymer micelles can be varied by introducing a diverse class of 

hydrotropic structures into the block copolymers through ATRP. 28,29,78–80 Table 19.3 lists some 

of the block and graft copolymers consisting of poly(ethylene glycol) (PEG) and the hydrotropic 

polymers that have been synthesized based on the molecular structures of identified hydrotropic 

agents for paclitaxel, such as DENA and PNA. 

CH3 OCH2CH2 n OH + Br 

O 

C 

N 

N 

O 

O 

Cu(I)Br, PMDETA, 80 °C 

O 

O 

N 

NH N 

CH3 CH Br 

TEA/DMAP 

CH 2Cl 2 

PEG−PDENA 

O 

O 

PEG−P(2-VBOPNA) 


O 

CH3 OCH2CH2 O C 

n 

CH 3 

CH3 OCH2CH2 O C CH CH2 CH Br 

n m 

CH3 

CH3 OCH2CH2 O C CH CH 

n 2 CH Br 

m 

N 

N 

O 

O 

O 

O 

N 

CH 3 

CH 

NH N 

FIGURE 19.5 Synthetic routes for PEG–PDENA and PEG–P(2-VBOPNA) block copolymers. 


Br


The main components necessary for the synthesis of the block copolymers are PEG modified 

with bromine or chlorine, and the modified hydrotropic agent with polymerizable vinyl, acryl, and 

styryl groups. Table 19.4 lists the useful hydrotropic agents modified with double bond functionality. 

The relevant combination of monomers listed in Table 19.4 can lead to a diverse class of the 

copolymers capable of making the hydrotropic micelles. 

In addition to the structural variation, a new property such as stimuli-responsibility can be 

endowed with the hydrotropic polymer micelles. One example is stimuli-responsive hydrotropic 

polymer micelles which were derived from PEG–P(2-VBOPNA) copolymers. In this system, PEG 

is freely water-soluble, independent of pH, whereas P(2-VBOPNA) is known to be soluble in water 

only when the hydrotropic PNA groups are protonated. P(2-VBOPNA) is soluble in low pH but 

loses its water-solubility above a critical pH (pH 2.5). Thus, the PEG–P(2-VBOPNA) copolymers, 

as expected, dissolved molecularly as unimers at low pH but become amphiphilic above pH 2.5. 

The combination of PEG and P(2-VBOPNA) in a block copolymer provides an opportunity to 

undergo pH-responsive micellization by controlling solution pH. Figure 19.6 shows pH-dependency 

of scattering intensity of PEG–P(2-VBOPNA) by dynamic light scattering. 

As solution pH increases, scattering intensity increases abruptly above pH 2.0, accompanied by 

the formation of micelles around pH 3.0 of which sizes were about 35 nm. As pH of the aqueous 

solution of the block copolymers is increased from 2 to 3, the degree of protonation of PNA groups 

decreased, and thus the P(2-VBOPNA) block becomes progressively more hydrophobic, as 

expected. 1 H NMR analysis is also a good tool to confirm the pH-responsive micellization of 

PEG–P(2-VBOPNA) (Figure 19.7). 

At pH 1, the block copolymers are fully solvated, and thus the signals assigned for the each 

block are clearly visible. On the other hand, at pH 7.4, the resonance peaks of P(2-VBOPNA) 

disappear while the signals from PEG are still evident, which indicates the poor solvation and/or 

limited molecular motion of P(2-VBOPNA) blocks. This result strongly supports the formation of 

micelles consisting of P(2-VBOPNA) as a hydrophobic inner core and PEG as a hydrated outer 

shell. Using this unique pH-sensitive micelle system, it may be suggested that PEG-b- 

P(2-VBOPNA) allows a unique drug-loading method into polymeric micelles, where drug 

molecules are initially solubilized exclusively by the interaction with hydrotropic 

P(2-VBOPNA) block at low pH, and then drugs are possibly encapsulated into inner core of 

micelles upon pH-responsive micellization. The use of hydrotropic block copolymers showing 

pH-responsive micellization does not need organic solvents during the drug loading process. 

The ideal oral delivery systems for anti-cancer drugs need the two following physicochemical 

properties to maximize oral bioavailability of drugs: the fast drug release during the transit of GI, 

and the prolonged retention time in the GI tract. The former was known to be the characteristic of 

hydrotropic polymer micelles. Thus, if the design of hydrotropic polymer micelles showing 

mucoadhesion property is possible, the idealized delivery systems can be developed based on 

hydrotropic polymer micelles. The study on the degradation kinetics of paclitaxel in aqueous 

TABLE 19.3 

Exemplary Copolymers of PEG and Hydrotropic Polymers Synthesized 

from Modified Hydrotropic Agents 

Poly(ethylene glycol)-block-poly(2-(4-vinylbenzyloxy)-N,N-diethylnicotinamide) 

Poly(ethylene glycol)-block-poly(2-(4-vinylbenzyloxy)-N-picolylnicotinamide) 

Poly(ethylene glycol)-block-poly(2-(4-vinylbenzyloxy)-nicotinamide) 

Poly(oligoethylene glycol methacrylate-co-poly(2-(4-vinylbenzyloxy)-N,N-diethylnicotinamide) 

Poly(oligoethylene glycol methacrylate-co-poly(2-(4-vinylbenzyloxy)-N-picolylnicotinamide) 

Poly(oligoethylene glycol methacrylate-co-poly(2-(4-vinylbenzyloxy)-nicotinamide) 


TABLE 19.4 

Modified Hydrotropic Monomers That Can Be Used in the Synthesis of Hydrotropic Polymer Micelles 

N 

CH 2 CH 

O O 

N 

HN 

O 

N 

CH2 CH 

O 

N 

HN 

3 

O 

O 

O 

6-acryloyl-N-picolylnicotinamide 6-(2-(acryloyl) ethoxyethoxyethoxy)-N-picolylnicotinamide 


N 

CH 2 CH 

N 

HN 

O 

O 

6-(4-vinylbenzyl oxy)-N-picolyl 

nicotinamide 

CH 2 CH 

O O 

N 

O 

N 

2-(2-(acryloyl) ethoxyethoxyethoxy)-N, Ndiethylnicotinamide 

CH 2 

O 

N 

CH 

3 O 

O 

O 

N 

2-(4-vinylbenzyloxy)-N, N-diethyl 

nicotinamide 

CH 2 CH 

N 

O 

O 

N 

2-acryloyl-N, N-diethylnicotinamide 

398 



Scattering intensity (Kcps) 

1400 

1200 

1000 

800 

600 

0 

1 

2 

3 

FIGURE 19.6 Variation of scattering intensity (C) as a function of pH for PEG 5000-P(2-VBOPNA) 2070 at 

5 mg/mL (nZ3). 

10 9 

(a) 

pH 1.0 

4 

pH 

d (PPM) 

10 9 8 7 6 5 4 3 2 1 0 

(b) 

8 

pH 7.4 

7 

6 

DOH 

5 

DOH 

d (PPM) 

FIGURE 19.7 1 H NMR spectra of PEG5000-P(2-VBOPNA)2070 at 20 mg/mL in D2O: (a) pH 1.0; (b) pH 7.4. 


4 

5 

3 

6 

2 

7 

1 

8 

0

400 


solutions at different pH values shows that a maximum stability of paclitaxel occurred in the pH 3–5 

region. 61 This provides a perfect opportunity to prepare mucoadhesive hydrotropic polymer micelle 

formulation for oral paclitaxel delivery. One of the best mucoadhesive polymers is 

poly(acrylic acid) (PAA), which is highly adhesion to biological mucosa at a pH lower than 5. 

PAA at neutral pH is highly charged and highly water-soluble, and can provide a hydrophilic 

segment of the block copolymers. Upon contact with the stomach’s low pH, the polymer 

becomes highly mucoadhesive. Alternatively, PAA can be treated at pH 5 to provide the mucoadhesive 

property. Because of the highly hydrophilic properties of PAA, it can replace the PEO 

segment of the block copolymers. Thus, the design of the block copolymer consisting of PAA 

and the hydrotropic block may provide the useful oral formulations of poorly soluble anticancer 

agents. 

19.4.5 HYDROTROPIC POLYMER MICELLES AS CARRIERS FOR ANTI-CANCER DRUGS 

Hydrotropic micelles were first examined for their ability to solubilize paclitaxel, a good example 

of an anti-cancer drug with extremely low water solubility. 27 Due to the high potential of paclitaxel, 

many formulations have been developed to increase water solubility. Although the polymeric 

micelles based on the amphiphilic block copolymers have shown high potential as drug solubilizing 

systems, most polymeric micelles have shown limited solubilizing capacity for paclitaxel, and, in 

most cases, maximum contents of paclitaxel loaded into micelles was around 20 wt%. 9,20,25 

Besides, a simple polymer design may not predict whether the resulting polymer micelles show 

high solubilizing capacity. A more serious limitation is the poor stability of paclitaxel-solubilized 

polymeric micelles in water, which becomes lower as the content of paclitaxel increases. 20 Of the 

properties of solubilizing systems, a high solubilizing capacity and a good physical stability are the 

two most important factors in determining whether the drug delivery systems are clinically useful. 

The goal of using hydrotropic polymer micelles is to develop a hydrotropic polymeric micelle that 

has a high solubilizing capacity for poorly water-soluble drugs, as well as a good long-term 

stability. Because an identified hydrotrope for a specific drug is introduced as a core component 

in a highly localized way, paclitaxel solubilization may be presented by the synergistic effect of 

both hydrotropic and micellar solubilization. The poor colloidal stability of existing micelles is 

normally caused by the enhanced hydrophobicity of micelles after solubilization of paclitaxel. 

Hydrotropic moieties, characterized by a strong hydrophilic nature, are expected to allow good 

stability for paclitaxel-loaded micelles in water. Huh et al. described the superiority of the hydrotropic 

polymer micelles relative to current polymeric micelles, and how the hydrotropic polymer 

micelles are different from existing systems. 27 These questions can be answered by comparing their 

drug loading and physical stability. The loading capacity of the normal polymer micelles for poorly 

soluble drugs is decided by various factors, such as the length of the core-forming polymer, and 

compatibility between drugs and the core-forming polymers. 1 Of these factors, the compatibility is 

the most significant in determining the solubilizing capacity. One parameter that has been used to 

assess the compatibility between solubilizates and the polymer is Flory–Huggins interaction parameter 

as described in Section 19.2.4. 1 This value is dependent on both the selected drug and the 

polymer. Due to the uniqueness of each drug, no single core-forming block can maximize the 

solubilization level for all drugs. Therefore, the first priority is to find or synthesize the right 

structure for effective solubilization of selected drugs. However, the number of biocompatible 

polymers is limited, and an added difficulty of the synthesis step is the screening of a large 

number of the polymer structures for effective solubilization of the selected drug. On the other 

hand, the hydrotropic approach is simpler and less labor-intensive, even though a screening process 

is also required. Many hydrotropes can be identified for a selected drug by a simple mixing 

procedure, and the broad range of the chemical structure can be readily screened. 25 The key 

concept of the hydrotropic polymer micelles is based on the hydrotrope containing core-forming 

polymers. Thus, the systematic design affords more efficient systems for solubilizing poorly soluble 



anti-cancer drugs. Another advantage of this approach is the enhanced physical stability of the 

formulations. This property is one measure that distinguishes hydrotropic polymer micelles from 

normal polymer micelles. Conventional polymer micelles have a polymer with a strong hydrophobic 

nature, specifically the drug solubilization has been expected only from the hydrophobic 

interaction between drugs and the inner core. Often, polymer micelles with drugs cannot overcome 

the enhanced hydrophobicity and the secondary aggregation between micelles, resulting in precipitation. 

From this point of view, the hydrotropic polymer micelles provide the formulations with 

good stability, even at a high loading of poorly soluble anti-cancer drugs, due to the hydrophilic 

nature of the hydrotropes residing in the micellar core domains. Of course, the same approach can 

be used for solubilization of other poorly soluble drugs. Huh et al. reported results on the solubilizing 

(loading) effect of the hydrotropic polymer micelle for paclitaxel by a dialysis method. 

The results are illustrated in Figure 19.8. 

The hydrotropic polymeric micelles solubilized paclitaxel at a level of 18.4–37.4 wt%, 

depending on the organic solvents used in the dialysis and the initial feed weight ratio of paclitaxel 

to the block copolymer. The loading content normally increases to certain feed-weight ratios of 

paclitaxel to the block copolymer. However, as the amount of paclitaxel was further increased, 

precipitates of unloaded paclitaxel formed during dialysis, resulting in the decreased loading 

contents. The maximum loading was observed with initial feed weight ratio of 1:0.5. When acetonitrile 

and the feed weight ratio of 1:0.5 were used the loading content was as high as 37.4 wt%, 

which was not possible with existing polymeric micelle systems. The maximum loading content of 

paclitaxel in a control micelle of PEG 2000–PDLLA 2000 was estimated to be 27.6 wt%, which is close 

to the literature value. Loading capacity of PEG–PDENA micelles for paclitaxel was enhanced by 

increasing the block length of PDENA. 

Poor physical stability of drug-loaded micelles, which causes serious problems in drug formulation, 

has been often addressed. There is often a compromise between loading content and stability 

in polymeric micelle systems. Drug-loaded micelles may break up more easily in aqueous media, 

resulting in drug precipitation. The physical stability of polymer micelles was investigated by 

several methods and compared. PEG–PLA micelles that have been extensively used to solubilize 

paclitaxel demonstrated a high loading capacity, but showed drug precipitation after 24 h. 

The initial transparent micelle solution became translucent, or turbid, after 24 h. Similar results 

have been reported in the literature. 20,21,25 On the other hand, PEG–PDENA micelle solutions were 

observed to be stable for several weeks without drug precipitation at the same or higher loading 

contents of paclitaxel. 

Paclitaxel content (wt%) 

50 

40 

30 

20 

10 

0 

Acetonitrile DMAc DMF 

1 : 0.25 

1 : 0.3 

1 : 0.35 

1 : 0.4 

1 : 0.5 

1 : 0.6 

Control 

FIGURE 19.8 Paclitaxel lading contents in hydrotropic polymer micelles and comparison with 

PEG2000–PDLLA2000 micelles. 


402 

The loading content of paclitaxel in polymer micelles was reported as a function of time. 

As shown in Figure 19.9, the paclitaxel content in PEG–PDENA micelles was maintained for 

more than 30 days. 27 

In contrast, the PEG–PLA micelle showed a dramatic decrease in paclitaxel content after three 

days, which resulted from drug precipitation. The stability of drug-loaded PEG–PDENA micelles 

was further confirmed by observing no change in micelle size and scattering intensity using 

dynamic light-scattering measurements. 

The paclitaxel release profiles of PEG–PDENA and PEG–PLA micelles were compared as 

shown in Figure 19.10. 27 

For PEG–PDENA micelles, the micelles of 25.9 wt% loading released the most drug within 

48 h, whereas the micelles of 31.3 wt% drug loading showed a much faster drug release that was 

complete within 24 h. It was found that the faster release was from the micelles with higher drug 

loading. The micelles with higher drug loading lead to a relatively lower polymer concentration. 

Thus, there is less polymer–drug interaction in the micelles, thereby resulting in faster release 

kinetics. The in vitro release data showed that paclitaxel was released faster from the hydrotropic 

polymer micelles than from PEG to PLA micelles. The PEG–PLA micelles required almost 72 h to 

release all the loaded paclitaxel. As indicated by the higher CMC values of PEG–PDENA micelles, 

the hydrotropic polymeric micelles have less hydrophobic cores that may release the loaded 

paclitaxel faster than more hydrophobic PLA cores. On the other hand, the release from paclitaxel 

bulk powder was negligible, even after 74 h. (Figure 19.10). 

The cytotoxicity of anti-cancer agents loaded in polymer micelle formulations is one measure 

for determining whether the formulation is effective in delivering the drugs into the tumor cells. 

The anti-tumor cytotoxicities of hydrotropic polymeric micelles and other control micelle systems 

on various cell lines, as measured by ED50, are shown in Table 19.5. The resulting cytotoxicity of 

hydrotropic polymer micelles and control micelles clearly show the superior cytotoxic properties 

of hydrotropic micelles. The ED50 values of hydrotropic micelles are much lower than PEG–PLA 

and PEG–PPA micelles. The most widely used polymeric micelles is PEG–PLA micelles, with the 

maximum paclitaxel loading capacity of 24 wt%. Although the hydrotropic polymer micelles have 

higher paclitaxel, loading up to 37 wt%, hydrotropic polymer micelles with only 20 and 25 wt% 

paclitaxel loading were used to compare their efficacy with that of PEG–PLA micelles. 

At equivalent paclitaxel loading, i.e., 25 wt% loading, hydrotropic polymer micelles were 

Paclitaxel concentration (%) 

120 

100 

80 

60 

40 

20 

14571g9 eps 

PEG−PDENA (25.9 wt% paclitaxel) 

PEG−PDENA (18.4 wt% paclitaxel) 

PEG−PLA (27.6 wt% paclitaxel) 

PEG−PLA (17.6 wt% paclitaxel) 

0 

0 5 10 15 

Time (day) 

20 25 30 


FIGURE 19.9 Changes in paclitaxel concentration of polymer micelles in distilled water. (From Huh, K.M., 

Lee, S.C., Cho, Y.W., Lee, J., Jeong, J.H., and Park, K., Journal of Controlled Release, 101, 59–68, 2005. With 

permission.) 



0.12 

0.08 

0.04 

0.00 

0 20 40 60 80 100 

Time (hour) 

FIGURE 19.10 Release kinetics from polymer micelles and paclitaxel powders in 0.8 M sodium salicylate at 

378C. The total amount of loaded paclitaxel was 0.1 mg. (From Huh, K. M., Lee, S. C., Cho, Y. W., Lee, J., 

Jeong and Park, K., Journal of Controlled Release, 101, 59–68, 2005. With permission.) 

substantially more effective. In the case of the MDA231 cell line, hydrotropic polymer micelles 

were more than two orders of magnitude more effective. The data clearly indicate that hydrotropic 

polymer micelles are not only more stable in aqueous solution, but also more effective for paclitaxel 

delivery to the cells. 

As mentioned in Section 19.4.2, the enhanced drug solubility can improve oral bioavailability 

of drugs by suppressing the MDR effect. The excellent solubilizing ability of hydrotropic polymer 

micelles for poorly soluble drugs can overcome this barrier, thereby enhancing drug absorption in 

the GI tract. As a preliminary study, PEG–PDENA micelles were examined for the bioavailability 

of loaded paclitaxel using an in vivo, chronically-catheterized rat model. Our initial study showed 

that oral delivery of paclitaxel using PEG–PDENA micelles could result in the blood paclitaxel 

TABLE 19.5 

ED 50 (mg/mL) of Paclitaxel and Paclitaxel (PTX)-Loaded Polymeric Micelles on Various 

Tumor Cell Lines 

Samples a 

HT-29 

Cancer Cell Lines 

MDA231 MCF-7 SKOV-3 

Doxorubicin (positive 

control) 

0.044 0.050 0.773 0.611 

Paclitaxel 0.003 0.033 0.043 0.006 

PTX-loaded HTM 

(25 wt% loading) 

0.005 0.002 0.002 0.001 

PTX-loaded HTM 


0.006 0.004 ! 0.001 0.008 

PTX-loaded PLA-PEG 


0.014 0.305 ! 0.001 0.015 

PTX-loaded PEG–PPA 0.012 1.077 0.002 1.543 

HTM alone 4.221 5.650 5.767 0.048 

PEG–PLA alone 4.672 8.435 5.533 — 

PEG–PPA alone 0.277 4.881 6.194 — 

a 

PTX, paclitaxel; HTM, hydrotropic micelle; PLA, poly(lactic acid); PEG, poly(ethylene glycol); PPA, poly(phenylala- 

nine). 


404 

concentrations that are clinically significant. To increase oral bioavailability, hydrotropic polymer 

micelles with other interesting functions can be developed. The systematic design of hydrotropic 

polymer micelles for its faster drug releasing properties, as well as the prolonged retention during 

through the GI tract, can provide an opportunity to develop ideal oral delivery systems for poorly 

soluble anti-cancer drugs. 

19.5 CONCLUSIONS AND FUTURE PERSPECTIVES 

Hydrotropic polymer micelle formulation presents an alternative and promising approach in formulation 

of poorly soluble drugs. Based on synergistic effect of the micellar characteristics and 

hydrotropic activity, the hydrotropic polymer micelles exhibit a high drug loading capacity with 

enhanced long-term stability in aqueous media. These unique properties make it highly attractive 

for applications of controlled delivery of poorly soluble anti-cancer drugs. The hydrotropic polymer 

micelles can be applied to many poorly soluble drugs and drug candidates by introducing the 

identified hydrotropic structures for specific drug molecules. The hydrotropic polymer micelle is 

in the early stages of the development, but due to its unique properties, the systematic design may 

generate effective oral formulations applicable to many poorly soluble anti-cancer agents. The high 

versatility of the hydrotropic polymer micelle in creating a broad range of drug formulations gives it 

the potential for a broad range of formulations in pharmaceutical and biomedical applications. 

REFERENCES 


1. Allen, C., Maysinger, D., and Eisenberg, A., Nanoengineering blocks copolymer aggregates for drug 

delivery, Colloids and Surfaces B: Biointerfaces, 16, 3–27, 1999. 

2. Lee, S. C., Kim, C., Kwon, I. C., Chung, H., and Jeong, S. Y., Polymeric micelles of poly(2-ethyl-2oxazoline)-block-poly(3-caprolactone) 

copolymer as a carrier for paclitaxel, Journal of Controlled 

Release, 89, 437–446, 2003. 

3. Yokoyama, M., Okano, T., Sakurai, Y., Fukushima, S., Okamoto, K., and Kataoka, K., Selective 

delivery of adriamycin to a solid tumor using a polymeric micelle carrier system, Journal of Drug 

Targeting, 7, 171–186, 1999. 

4. Lee, S. C., Chang, Y., Yoon, J.-S., Kim, C., Kwon, I. C., Kim, Y.-H., and Jeong, S. Y., Synthesis and 

micellar characterization of amphiphilic diblock copolymers based on poly(2-ethyl-2-oxazoline) and 

aliphatic polyesters, Macromolecules, 32, 1847–1852, 1999. 

5. Kwon, G., Naito, M., Yokoyama, M., Okano, T., Sakurai, Y., and Kataoka, K., Micelles based on AB 

block copolymers of poly(ethylene oxide) and poly(beta-benzyl L-aspartate), Langmuir, 9, 945–949, 

1993. 

6. Bader, H., Ringsdorf, H., and Schmidt, B., Watersoluble polymers in medicine, Die Angewandte 

Makromolekulare Chemie, 123/124, 457–485, 1984. 

7. Jeong, Y.-I., Cheon, J.-B., Kim, S.-H., Nah, J.-W., Lee, Y.-M., Sung, Y.-K., Akaike, T., and Cho, 

C. S., Clonazepam release from core-shell type nanoparticles in vitro, Journal of Controlled Release, 

51, 169–178, 1998. 

8. Kim, I. S. and Kim, S. H., Development of a polymeric nanoparticulate drug delivery system: in vitro 

characterization of nanoparticles based on sugar-containing conjugates, International Journal of 


9. Kim, S. Y. and Lee, Y. M., Taxol-loaded block copolymer nanospheres composed of methoxy 

poly(ethylene glycol) and poly(3-caprolactone) as novel anticancer drug carriers, Biomaterials, 

22, 1697–1704, 2001. 

10. Kwon, G. S. and Okano, T., Polymeric micelles as new drug carriers, Advanced Drug Delivery 

Review, 21, 107–116, 1996. 

11. Creutz, S., van Stam, J., De Schryver, F. C., and Jerome, R., Dynamics of poly((dimethylamino)alkyl 

methacrylate-block-sodium methacrylate) micelles. Influence of hydrophobicity and molecular 

architecture on the exchange rate of copolymer molecules, Macromolecules, 31, 681–689, 1998. 



12. Zhang, Y., Jin, T., and Zhuo, R., Methotrexate-loaded biodegradable polymeric micelles: Preparation, 

physicochemical properties and in vitro drug release, Colloids and Surfaces B. Biointerfaces. 

44, 104–109, 2005. 



micelles and their design for in vivo delivery to a solid tumor, Journal of Controlled Release, 50, 

79–92, 1998. 

14. Yokoyama, M., Opanasopit, P., Okano, T., Kawano, K., and Maitani, Y., Polymer design and incorporation 

methods for polymeric micelle carrier system containing water-insoluble anti-cancer agent 

camptothecin, Journal of Drug Targeting, 12, 373–384, 2004. 

15. You, L.-C., Lu, F.-Z., Li, Z.-C., Zhang, W., and Li, F. -M, Glucose-sensitive aggregates formed by 

poly(ethylene oxide)-block-poly(2-glucosyloxyethyl acrylate) with concanavalin A in dilute aqueous 

medium, Macromolecules, 36, 1–4, 2003. 

16. Harada, A. and Kataoka, K., Formation of polyion complex micelles in an aqueous milieu from a pair 

of oppositrly-charged block copolymers with poly(ethylene glycol) segments, Macromolecules, 28, 

5294–5299, 1995. 

17. Harada, A. and Kataoka, K., Novel polyion complex micelles entrapping enzyme molecules in the 

core: Preparation of narrowly-distributed micelles from lysozyme and poly(ethylene glycol)–poly 

(aspartic acid) block copolymers in aqueous medium, Macromolecules, 31, 288–294, 1998. 

18. Suh, H., Jeong, B., Rathi, R., and Kim, S. W., Regulation of smooth muscle cell proliferation using 

paclitaxel-loaded poly(ethylene oxide)–poly(lactide/glycolide) nanospheres, Journal of Biomedical 

Research Materials, 42, 331–338, 1998. 

19. Soga, O., van Nostrum, C. F.., Fens, M., Rijcken, C. J. F., Schiffelers, R. M., Storm, G., and Hennink, 

W. E., Thermosensitive and biodegradable polymeric micelles for paclitaxel delivery, Journal of 


20. Burt, H. M., Zhang, X., Toleikis, P., Embree, L., and Hunter, W. L., Development of copolymers of 

poly(D,L-lactide) and methoxypolyethylene glycol as micellar carriers of paclitaxel, Colloids and 



micellar carriers of taxol, International Journal of Pharmaceutics, 132, 195–206, 1996. 

22. Balasubramanian, D., Srinivas, V., Gaikar, V. G., and Sharma, M. M., Aggregation behavior of 

hydrotropic compounds in aqueous solution, Journal of Physical Chemistry, 93, 3865–3870, 1989. 

23. Srinivas, V. and Balasubramanian, D., Proline is a protein-compatible hydrotrope, Langmuir, 

11, 2830–2833, 1995. 

24. Srinivas, V. and Balasubramanian, D., When does the switch from hydrotropy to micellar behavior 

occur?, Langmuir, 14, 6658–6661, 1998. 

25. Lee, J., Lee, S. C., Acharya, G., Chang, C.-J., and Park, K., Hydrotropic solubilization of paclitaxel: 

Analysis of chemical structures for hydrotropic property, Pharmaceutical Research, 20, 1022–1030, 

2003. 

26. Lee, S. C., Acharya, G., Lee, J., and Park, K., Hydrotropic polymers: Synthesis and characterization of 

polymers containing picolylnicotinamide moieties, Macromolecules, 36, 2248–2255, 2003. 

27. Huh, K. M., Lee, S. C., Cho, Y. W., Lee, J., Jeong, J. H., and Park, K., Hydrotropic polymer micelle 

system for delivery of paclitaxel, Journal of Controlled Release, 101, 59–68, 2005. 

28. Gadelle, F., Koros, W. J., and Schechter, R. S., Solubilization of aromatic solutes in block copolymers, 


29. Gan, Z., Jim, T. F., Li, M., Yuer, Z., Wang, S., and Wu, C., Enzymatic biodegradation of poly 

(ethylene oxide-b-3-caprolactone) diblock copolymer and its potential biomedical applications, 


30. Alexandridis, P., Holzwarth, J. F., and Hatton, T. A., Micellization of poly(ethylene oxide)–poly 

(propylene oxide)–poly(ethylene oxide) triblock copolymers in aqueous solutions: Thermodynamics 

of copolymer association, Macromolecules, 27, 2414–2425, 1994. 

31. Kabanov, A. V., Slepnev, V. I., Kuzetsova, L. E., Batrakova, E. V., Alakhov, V. Y., Melik-Nubarov, 

N. S., Sveshinikov, P. G., and Kabanov, V. A., Pluronic micelles as tool for low-molecular compound 

vector delivery into a cell: Effect of Staphylococcus aureus enterotoxin B on cell loading with micelle 

incorporated fluorescent dye, Biochemistry International, 26, 1035–1042, 1992. 


406 



polymeric drug indomethacin-incorporated poly(ethylene oxide)–poly(benzyl-2 0 -aspartate) block 

copolymer micelles, Journal of Pharmaceutical Sciences, 85, 85–90, 1996. 

33. Kalyanasundaram, K. and Thomas, J. K., Environmental effects on vibronic band intensities in pyrene 

monomer fluorescence and their application in studies of micellar systems, Journal of the American 

Chemical Society, 99, 2039–2044, 1977. 

34. Jeong, B., Bae, Y. H., and Kim, S. W., Biodegradable thermosensitive micelles of PEG–PLGA–PEG 

triblock copolymers, Colloids and Surfaces B: Biointerfaces, 16, 185–193, 1999. 

35. Huh, K. M., Lee, K. Y., Kwon, I. C., Kim, Y.-H., Kim, C., and Jeong, S. Y., Synthesis of triarmed 

poly(ethylene oxide)–deoxycholic acid conjugate and its micellar characteristics, Langmuir, 16, 

10566–10568, 2000. 

36. Kabanov, A. V., Nazarova, I. R., Astafieva, I. V., Batrakova, E. V., Alakhov, V. Y., Yaroslavov, A. A., 

and Kabanov, V. A., Micelle formation and solubilization of fluorescent probes in poly(oxyethyleneb-oxypropylene-b-oxyethylene) 


37. Lee, K. Y., Jo, W. H., Kwon, I. C., Kim, Y.-H., and Jeong, S. Y., Structural determination and interior 

polarity of self-aggregates prepared from deoxycholic acid-modified chitosan in water, Macromolecules, 

31, 378–383, 1998. 

38. Kim, C., Lee, S. C., Kwon, I. C., Chung, H., and Jeong, S. Y., Complexation of poly(2-ethyl-2oxazoline)-block-poly(3-caprolactone) 

micelles with multifunctional carboxylic acids, Macromolecules, 

35, 193–200, 2002. 

39. Xu, R., Winnik, M. A., Hallett, F. R., Riess, G., and Croucher, M. D., Light-scattering study of the 

association behavior of styrene–ethylene oxide block copolymers in aqueous solution, Macromolecules, 

24, 87–93, 1991. 

40. Ringsdorf, H., Venzmer, J., and Winnik, F. M., Fluorescence studies of hydrophobically modified 

poly(N-isopropylacrylamide), Macromolecules, 24, 1678–1686, 1991. 

41. Wilhelm, M., Zhao, C.-L., Wang, Y., Xu, R., and Winnik, M. A., Poly(styrene–ethylene oxide) block 

copolymer micelle formation in water: a fluorescence probe study, Macromolecules, 24, 1033–1040, 

1991. 

42. Zhang, L. and Eisenberg, A., Multiple morphologies and characteristics of “crew-cut” micelle-like 

aggregates of polystyrene-b-poly(acrylic acid) diblock copolymers in aqueous solutions, Journal of 

the American Chemical Society, 118, 3168–3181, 1996. 

43. Zhang, L. and Eisenberg, A., Multiple morphologies of “Crew-Cut” aggregates of polystyrene-bpoly(acrylic 

acid) block copolymers, Science, 268, 1728–1731, 1995. 

44. Zhang, L., Yu, K., and Eisenberg, A., Ion-induced morphological changes in “Crew-Cut” aggregates 

of amphiphilic block copolymers, Science, 272, 1777–1779, 1996. 

45. Heise, A., Hedrick, J. L., Frank, C. W., and Miller, R. D., Starlike block copolymers with amphiphilic 

arms as models for unimolecular micelles, Journal of the American Chemical Society, 121, 

8647–8648, 1999. 

46. Newkome, G. R., Moorefield, C. N., Baker, G. R., Saunders, M. J., and Grossman, S. H., Unimolecular 

micelles, Angewandte Chemie International Edition in English, 30, 1178–1180, 1991. 

47. Zhang, G.-D., Harada, A., Nishiyama, N., Jeang, D.-L., Koyama, H., Aida, T., and Kataoka, K., 

Polyion complex micelles entrapping cationic dendrimer porphyrin: Effective photosensitizer for 

photodynamic therapy of cancer, Journal of Controlled Release, 93, 141–150, 2003. 

48. Jones, M.-C., Ranger, M., and Leroux, J.-C., pH-Sensitive unimolecular polymeric micelles: 

Synthesis of a novel drug carrier, Bioconjugate Chemistry, 14, 774–781, 2003. 

49. Liu, H., Farrell, S., and Uhrich, K., Drug release characteristics of unimolecular polymeric micelles, 


50. Nagasaki, Y., Okada, T., Scholz, C., Iijima, M., Kato, M., and Kataoka, K., The reactive polymeric 

micelles based on an aldehyde-ended poly(ethylene glycol)/poly(lactide) block copolymer, Macromolecules, 

31, 1473–1479, 1998. 

51. Torchilin, V. P., Structure and design of polymeric surfactant-based drug delivery systems, Journal of 


52. Le Garrec, D., Gori, S., Luo, L., Lessard, D., Smith, D. C., Yessine, M.-A., Ranger, M., and Leroux, J.- 

C., Poly(N-vinylpyrrolidone)-block-poly(D,L-lactide) as a new polymeric solubilizer for hydrophobic 

anticancer drugs: in vitro and in vivo evaluation, Journal of Controlled Release, 99, 83–101, 2004. 



53. Liggins, R. T. and Burt, H. M., Polyether–polyester diblock copolymers for the preparation of 

paclitaxel loaded polymeric micelle formulations, Advanced Drug Delivery Reviews, 54, 191–202, 

2002. 

54. Liu, S. Q., Tong, Y. W., and Yang, Y. Y., Incorporation and in vitro release of doxorubicin in 

thermally sensitive micelles made from poly(N-isopropylacrylamide-co-N,N-dimethylacrylamide)b-poly(D,L-lactide-co-glycolide) 

with varying compositions, Biomaterials, 26, 5064–5074, 2005. 

55. Lukyanov, A. N. and Torchilin, V. P., Micelles from lipid derivatives of water-soluble polymers as 

delivery systems for poorly soluble drugs, Advanced Drug Delivery Reviews, 56, 1273–1289, 2004. 

56. U.S. National Institutes of Health. National Cancer Institute, 1986. Developmental Therapeutics 

Program. NCI Investigational Drugs: Pharmaceutical Data, pp. 70–72, (http://dtp.nci.nih.gov/docs/ 

idrugs/chembook.html/). 

57. Adams, J. D., Flora, K. P., Goldspiel, B. R., Wilson, J. W., and Arbuck, S. G., Taxol: a history of 

pharmaceutical development and current pharmaceutical concerns, Journal of the National Cancer 

Institute Monogram, 15, 141–147, 1993. 


long-circulating polymeric nanospheres, Science, 263, 1600–1603, 1994. 

59. Mazzo, D. J., Nguyen-Huu, J.-J., Pagniez, S., and Denis, P., Compatibility of docetaxel and paclitaxel 

in intravenous solutions with poyvinyl chloride infusion materials, American Journal of Health- 

System Pharmacy, 54, 566–569, 1997. 

60. Alkan-Onyuksel, H., Ramakrishnan, S., Chai, H.-B., and Pezzuto, J. M., A mixed micellar formulation 

suitable for the parenteral administration of taxol, Pharmaceutical Research, 11, 206–212, 

1994. 

61. Dordunoo, S. K. and Burt, H. M., Solubility and stability of taxol: Effects of buffer and cyclodextrins, 

International Journal of Pharmaceutics, 133, 191–201, 1996. 

62. Muller, R. H. and Bohm, B., Nanosuspensions, In Emulsions and Nanosuspensions for the Formulation 

of Poorly Soluble Drugs, Muller, Rainer H., Benita, S., and Bohm, B., Eds., Medpharm 

Scientific Publishers, Stuttgart, pp. 149–174, 1998. 

63. Rubino, J. T. and Yalkowsky, S. H., Cosolvency and cosolvent polarity, Pharmaceutical Research,4, 

220–230, 1987. 

64. Serajuddin, A. T. M., Solid dispersion of poorly water-soluble drugs: Early promises, subsequent 

problems, and recent breakthroughs, Journal of Pharmaceutical Science., 88, 1058–1066, 1999. 

65. Sharma, A. and Straubinger, R. M., Novel taxol formulations: Preparation and characterization of 

taxol-containing liposomes, Pharmaceutical Research, 11, 889–896, 1994. 

66. Tarr, B. D., Sambandan, T. G., and Yalkowsky, S. H., A new parenteral emulsion for the administration 

of taxol, Pharmaceutical Research, 4, 162–165, 1987. 

67. Nishiyama, N., Bae, Y., Miyata, K., Fukumura, S., and Kataoka, K., Drug Discovery Today: 

Technologies, 2, 21–26, 2005. 




antitumor efficacy, Bioconjugate Chemistry, 16, 122–130, 2005. 

69. Jaracz, S., Chen, J., Kuznetsova, L. V., and Ojima, I., Recent advances in tumor–targeting anticancer 

drug conjugates, Bioorganic & Medicinal Chemistry, 13, 5043–5054, 2005. 

70. van Nostrum, C. F., Polymeric micelles to deliver photosensitizers for photodynamic therapy, 


71. Torchilin, V. P., Polymeric micelles in diagnostic imaging, Colloids and Surfaces B: Biointerfaces, 

16, 305–319, 1999. 

72. Dolmans, D. and Fukumura, D., Photodynamic therapy for cancer, Nature Reviews Cancer, 3, 

380–387, 2003. 

73. Yalkowsky, S. H., Solubility and Solubilization in Aqueous Media, American Chemical Society, 

Washington, DC, 1999. 

74. Gandhi, N. N., Kumar, M. D., and Sathyamurthy, N., Effect of hydrotropes on solubility 

and mass-transfer coefficient of butyl acetate, Journal of Chemical & Engineering Data, 43, 

695–699, 1998. 


408 


75. Lobenberg, R., Amidon, G. L., and Vierira, M., Solubility as a limiting factor to drug absorption, In 

Oral Drug Absorption: Prediction and Assessment, Dressman, J. H. and Lennernas, H., Eds., Marcel 

Dekker, New York, pp. 137–153, 2000. 

76. Rasool, A. A., Hussain, A. A., and Dittert, L. W., Solubility enhancement of some water-insoluble 

drugs in the presence of nicotinamide and related compounds, Journal of Pharmaceutical Science, 80, 

387–393, 1991. 

77. Coffman, R. E. and Kildsig, D. O., Hydrotropic solubilization-Mechanistic studies, Pharmacal 

Research, 13, 1460–1463, 1996. 

78. Xia, J., Zhang, X., and Matyjaszewski, K., Atom transfer radical polymerization of 4-vinylpyridine, 


79. Matyjaszewski, K. and Xia, J., Atom transfer radical polymerization, Chemical Reviews, 101, 

2921–2990, 2001. 

80. Patten, T. E., Xia, J., Abernathy, T., and Matyjaszewski, K., Polymers with very low polydispersities 

from atom transfer radical polymerization, Science, 272, 866–868, 1996. 


20 

CONTENTS 

Tumor-Targeted Delivery of 

Sparingly-Soluble Anti-Cancer 

Drugs with Polymeric Lipid-Core 

Immunomicelles 


20.1 Introduction ....................................................................................................................... 409 

20.2 Micelles, Micellization, and Drug Delivery ..................................................................... 410 

20.3 Tumor Targeting of Polymeric Micelles .......................................................................... 411 

References..................................................................................................................................... 417 


Various drug delivery and drug targeting systems-based soluble polymers, nanoparticles made of 

natural and synthetic polymers, nanocapsules, lipoproteins, liposomes, micelles, and many other 

pharmaceutical carriers are currently widely used to minimize drug degradation and loss upon 

administration, prevent harmful or undesirable side-effects, and increase drug bioavailability and 

the fraction of the drug accumulated in the required (pathological) zone. 1,2 Many of those drug 

carriers are long-circulating 3,4 because prolonged circulation maintains the required therapeutic 

level of pharmaceuticals in the blood for extended time intervals, and it allows for slow accumulation 

of high-molecular-weight drugs or drug-containing nanocarriers in pathological sites with 

affected and leaky vasculature via the enhanced permeability and retention effect (EPR). 5,6 It also 

provides better targeting effects for specific ligand-modified drugs and drug carriers. 7 

The development of biocompatible, biodegradable, long-circulating, and specifically targeted 

drug carriers for sparingly soluble pharmaceuticals that possess high loading capacities is attracting 

a lot of attention because the therapeutic application of hydrophobic, poorly water-soluble agents 

might be associated with serious problems as low water-solubility results in poor absorption and 

low bioavailability upon oral administration. 8 The aggregation of poorly soluble drugs upon intravenous 

administration might lead to an embolism 9 and increased local toxicity at the sites of 

aggregate deposition. 10 However, the hydrophobicity (low water solubility) allows a drug molecule 

to better penetrate a cell membrane when the molecular target for the drug is intracellularly 

located, 11,12 and this is especially important for certain anti-cancer agents, many of which are 

bulky polycyclic compounds. 13 To overcome some drugs’ poor solubility and to increase their 


409

410 

bioavailability, certain clinically acceptable organic solvents are used in pharmaceutical formulations 

such as ethanol or polyethoxylated castor oil (Cremophor EL) 10 as well as drug carriers such 

as liposomes 14 and cyclodextrins. 15 In addition, a lot of interest is currently expressed toward using 

various micelle-forming substances in formulations of insoluble drugs. 10 

20.2 MICELLES, MICELLIZATION, AND DRUG DELIVERY 


Micelles represent dispersed systems, consisting of particulate matter or dispersed phase that are 

distributed within a continuous phase or dispersion medium and, under certain conditions (concentration 

and temperature), are spontaneously formed by amphiphilic or surface-active agents 

(surfactants) that are molecules that consist of two clearly distinct moieties differing in their affinity 

toward a given solvent. 16 They typically have particle sizes within a 5 to 50–100 nm range. The 

concentration of a monomeric amphiphile where micelles (aggregates, including several dozens of 

amphiphilic molecules and usually spherical in shape) appear is called the critical micelle (or 

micellization) concentration (CMC). The lower the CMC value of a given amphiphilic polymer, 

the more stable micelles are even at low net concentration of an amphiphile in the medium. This is 

important from the practical point of view because upon dilution with the large volume of the blood, 

only micelles with low CMC value still exist, whereas micelles with high CMC value may dissociate 

into uinimers, and their content may precipitate in the blood. In aqueous media, 

hydrophobic fragments of amphiphilic molecules form the core of a micelle that can solubilize 

poorly soluble pharmaceuticals, 17 whereas polar molecules will be adsorbed on the micelle 

surface, and substances with intermediate polarity/solubility will be distributed in various 

intermediate positions. 

Micellization of biologically active substances is a wide-spread phenomenon that is important 

for pharmacological action. For example, the increase in the bioavailability of a lipophilic drug 

upon oral administration is attributed to drug solubilization in the gut by naturally occurring biliary 

lipid/fatty acid-containing mixed micelles produced by organism upon the digestion of the dietary 

fat. Surfactant micelles are also widely used as adjuvants and drug carrier systems, and in many 

areas of pharmaceutical technology and controlled drug delivery research, micelles as drug carriers 

provide a set of clear advantages. 18–22 The solubilization of sparingly soluble drugs by micelleforming 

surfactants results in an increased water solubility of sparingly soluble drug and its 

improved bioavailability, extended blood half-life upon intravenous administration, reduction of 

toxicity and other adverse effects, enhanced permeability across the physiological barriers, and 

substantial changes in drug biodistribution. In addition, being in a micellar container, the drug is 

well-protected from possible inactivation under the effect of biological surroundings and does not 

provoke toxic side effects on non-target tissues. 

Polymeric micelles, especially popular in drug delivery research, are formed from block-copolymers 

consisting of hydrophilic and hydrophobic monomer units with the length of a hydrophilic 

block exceeding, to some extent, that of a hydrophobic one. These micelles contain the core of the 

hydrophobic blocks stabilized by the corona of hydrophilic polymeric chains. If the length of a 

hydrophilic block is too high, copolymers exist in water as unimers (individual molecules), whereas 

molecules with very long hydrophobic block forms structure with non-micellar morphology. 23 The 

core compartment of the pharmaceutical polymeric micelle should demonstrate high loading 

capacity, controlled release profile for the incorporated drug, and good compatibility between 

the core-forming block and incorporated drug. The micelle corona should provide an effective 

steric protection for the micelle. It should also determine for the micelle hydrophilicity, charge, 

the length and surface density of hydrophilic blocks, and the presence of reactive groups suitable 

for further micelle derivatization such as an attachment of targeting moieties. The listed properties 

control key biological characteristics of a micellar carrier including pharmacokinetics, 


Tumor-Targeted Delivery of Sparingly-Soluble Anti-Cancer Drugs 411 

biodistribution, biocompatibility, longevity, surface adsorption of biomacromolecules, adhesion to 

biosurfaces, and targetability. 1,24–30 

PEG still remains the hydrophilic block of choice, although some other hydrophilic polymers 

may be used to make corona blocks, 31,32 whereas a variety of polymers was used to build hydrophobic 

core-forming blocks including propylene oxide, 26,33 L-lysine, 34,35 aspartic acid, 36,37 

b-benzoyl-L-aspartate, 38,39 caprolactone, 40,41 D,L-lactic acid, 27,42 and many others. Phospholipid 

residues can also be successfully used as hydrophobic core-forming groups. 43 The use of lipid 

moieties as hydrophobic blocks capping hydrophilic polymer (such as PEG) chains can provide 

additional advantages for particle stability when compared with conventional amphiphilic polymer 

micelles because of the existence of two fatty acid acyls that may considerably contribute to an 

increase in the hydrophobic interactions in the micelle’s core. Although diacyllipid–PEG conjugates 

have been introduced into the area of controlled drug delivery as polymeric surface modifiers 

for liposomes, 44 the diacyllipid–PEG molecule itself represents a typical amphiphilic copolymer 

with a bulky hydrophilic (PEG) portion and a very short, but extremely hydrophobic, diacyllipid 

part. Similar to other PEG-containing amphiphilic block-copolymers, diacyllipid–PEG conjugates 

were found to form micelles in an aqueous environment. 45 A series of PEG–phosphatidylethanolamine 

(PEG–PE) conjugates was synthesized using egg PE and N-hydroxysuccinimide esters of 

methoxy–PEG succinates with different MW of PEG fragments. 44 HPLC-based gel permeation 

chromatography showed that these polymers formed micelles of different sizes in water. The 

stability of the polymeric micelles was confirmed by the fact that no dissociation into individual 

polymeric chains was found following the chromatography of the serially diluted samples of 

PEG–PE up to a polymer concentration of ca. 1 ml/ml corresponds to a micromolar CMC value 

that is at least 100-fold lower than those of conventional detergents. 46 Such low CMC values 

indicate that micelles prepared from these polymers will maintain their integrity even upon 

strong dilution (for example, in the blood during a therapeutic application). High stability of 

polymeric micelles allows also for good retention of encapsulated drugs in the solubilized form 

upon parenteral administration. PEG–PE micelles also can efficiently incorporate a broad variety of 

sparingly soluble and amphiphilic substances including paclitaxel, tamoxifen, camptothecin, 

porphyrine, vitamin K3, and others. 47–49 

Among other factors influencing the efficacy of drug loading into the micelle are the sizes of 

both core-forming and corona-forming blocks. 50 In the first case, the larger the hydrophobic block, 

the bigger core size and the ability of micelle to entrap hydrophobic drugs. In the second case, the 

increase in the length of the hydrophilic block results in the increase of the CMC value, i.e., at a 

given concentration of the amphiphilic polymer in solution the smaller fraction of this polymer will 

be present in the micellar form and the quantity of the micelle-associated drug drops. Incorporation 

into polymeric micelles is beneficial for various sparingly soluble drugs. Therefore, doxorubicin 

incorporated into Pluronic w micelles demonstrated superior properties as compared to free drugs in 

the experimental treatment of murine tumors (leukemia P388, myeloma, and Lewis lung carcinoma) 

and human tumors (breast carcinoma MCF-7) in mice. 51 In addition, the reduction of the side 

effects of the drug was observed in many cases. The toxicity decrease (smaller vascular damage and 

liver focal necrosis) was found for the polymeric micelle-incorporated anti-tumor drug KRN5500. 52 

20.3 TUMOR TARGETING OF POLYMERIC MICELLES 

Making micelles capable of specific accumulation in desired areas in the body (tumors) can further 

increase the efficiency of micelle-encapsulated pharmaceuticals. The simplest way to achieve this is 

to rely on previously mentioned preferential accumulation of drug-loaded micelles in areas with 

compromised vasculature (tumors and infarcts) via the EPR effect that is based on the spontaneous 

penetration of long-circulating macromolecules, molecular aggregates, and particulate drug carriers 

into the interstitium through the leaky blood vessels in certain pathological sites in the body. 


412 

It has been shown that the EPR effect is typical for solid tumors and infarcts. 5,6 Micelles prepared 

from PEG–PE conjugates with different length of the PEG block demonstrated much higher 

accumulation in tumors compared to non-target tissue (muscle) in experimental Lewis lung carcinoma 

(LLC: tumor with a relatively low vascular permeability) 53,54 in mice. The peak uptake is 

achieved around five hours post-injection, and the maximal total tumor uptake of the injected dose 

within the observation period (AUC) was found for micelles made PEG–PE with the largest size of 

the PEG block (5,000 Da). This was explained by the fact that these micelles had little extravasation 

into the normal tissue compared to micelles prepared from the shorter PEG–PE conjugates. 

Micelles prepared from PEG–PE conjugates with shorter PEG, however, might be more efficient 

carriers of poorly soluble drugs because they have a greater hydrophobic-to-hydrophilic phase ratio 

and can be loaded with drugs more efficiently on a weight-to-weight basis. Similar results were also 

obtained in another murine tumor model, EL4 T cell lymphoma (EL4); 55 see Figure 20.1. 

It is important that tumor diffusion and accumulation parameters of nanosized pharmaceutical 

carriers strongly depend on the cutoff size of tumor blood vessel wall, and the cutoff size varies for 

different tumors. 56–58 Because tumor vasculature permeability depends on the particular type of the 

tumor, 57 the use of micelles as drug carriers could be specifically justified for tumors whose 

vasculature has the low cutoff size (below 200 nm). The use of PEG–PE micelles for the effective 

delivery of a model protein drug (soybean trypsin inhibitor or STI, MW 21.5 kDa) to subcutaneously 

established LLC in mice was reported. 54 

An interesting targeting mechanism is based on the fact that many pathological processes in 

various tissues and organs are accompanied with local temperature increase and/or acidosis. 59,60 

Micelles made capable of disintegration under the increased temperature or decreased pH values in 

pathological sites (such as tumors) can combine the EPR effect-mediated accumulation with 

stimuli-responsiveness. It was shown that micelles made of thermo- or pH-sensitive components 

such as poly(N-isopropylacrylamide) and its copolymers with poly(D,L-lactide) and other blocks 

acquire the ability to disintegrate in target areas, releasing the micelle-incorporated drug. 61–65 

pH-responsive polymeric micelles loaded with phtalocyanine seem to be promising carriers for 

the photodynamic cancer therapy, 66 whereas doxorubicin-loaded polymeric micelles containing 

acid-cleavable linkages provided an enhanced intracellular drug delivery into tumor cells and, 

therefore, higher efficiency. 67 

As with many other delivery systems, the drug delivery potential of polymeric micelles may be 

further enhanced by attaching targeting ligands to their surface. The attachment of various specific 

ligand to the water-exposed termini of hydrophilic blocks has been used to improve the targeting 

of micelles and micelle-incorporated drugs and DNA. 26,68,69 Among those ligands, various sugar 

moieties can be named as well as 70 transferrin, 71 and folate residues 72 because many target cells, 

AUC dose/g organ 

75 

50 

25 

Tumor 

Muscle 

Tumor 

Muscle 

0 

0 

(a) PEG750-PE PEG2000-PE (b) PEG750-PE PEG2000-PE FIGURE 20.1 Accumulation of PEG–PE micelles in tumors in mice compared to muscle accumulation. 

(a) LLC tumor; (b) EL4 tumor. 


AUC dose/g organ 

75 

50 

25 



especially cancer cells, over express appropriate receptors (such as transferrin and folate receptors) 

on their surface. Transferrin-modified micelles based on PEG and polyethyleneimine with sizes 

between 70 and 100 are expected to target tumors with over expressed transferrin receptors. 71 

Similar targeting approaches were successfully tested with folate-modified micelles. 73 Poly(Lhistidine)/PEG 

and poly(L-lactic acid)/PEG block copolymer micelles carrying folate residue on 

their surfaces were shown to be efficient for the delivery of adriamycin to tumor cells in vitro, 

demonstrating potential for solid tumor treatment and combined targetability and pH-sensitivity. 74 

Among all specific ligands, monoclonal antibodies, primarily of the IgG class, and their Fab 

fragments provide the broadest opportunities in terms of diversity of targets and specificity of 

interaction. Several successful attempts to covalently attach an antibody to a surfactant or polymeric 

micelles (i.e., to prepare immunomicelles) have been already described. 22,26,71,75 By adapting 

the coupling technique developed for attaching specific ligands to liposomes utilizing PEG–PE with 

the free PEG terminus activated with p-nitrophenylcarbonyl (pNP) group, 76 PEG–PE-based immunomicelles 

have been prepared modified with monoclonal antibodies (see the principal reaction 

scheme in Figure 20.2). The micelle-attached protein was quantified using fluorescent labels or by 

SDS-PAGE, 75,77 and it was calculated that 10–20 antibody molecules could be attached to a single 

micelle. Antibodies attached to the micelle corona preserve their specific binding ability, and 

immunomicelles specifically recognize their target substrates as was confirmed by ELISA experiments 

while the micelles maintain essentially the same size as non-modified ones. 75 

One has to keep in mind, however, that whole antibody attachment to the micelles (as well as to 

any other drug carriers) might provoke faster clearance from the circulation as a result of uptake by 

Fc receptor-bearing Kupffer cells. To test if the antibody attachment affects the blood clearance of 

PEG–PE micelles, the comparison was made of the clearance characteristics of plain and antibodymodified 

micelles in mice; it was found that PEG–PE micelle modification with an antibody had a 

very small effect on their blood clearance. 75 

To specifically enhance the tumor accumulation of PEG–PE-based micelles, they have been 

modified with tumor-specific monoclonal antibodies. 75,77 Non-pathogenic monoclonal antinuclear 

autoantibodies with the nucleosome-restricted specificity (monoclonal antibody 2C5, mAb 2C5 

being among them) that recognize the surfaces of numerous tumor, but not normal, cells via tumor 

cell surface-bound nucleosomes 78,79 were used in these experiments. Because these antibodies bind 

a broad variety of cancer cells, they may serve as specific ligands for the delivery of drugs and drug 

carriers into tumors. The data shown in Figure 20.3a indicate that fluorescence (rhodamine)-labeled 

2C5-immunomicelles effectively bind to the surface of several unrelated human and murine tumor 

cells lines: human BT20 (breast adenocarcinoma), murine LLC (Lewis lung carcinoma), and EL4 

(T lymphoma) cells. The incubation of antibody-free micelles with the same cells results in 

virtually no micelle-to-cell association. Drug(paclitaxel)-loaded 2C5-immunomicelles also effectively 

bound various tumors cells (Figure 20.3b). Such specific recognition of cancer cells by drugloaded 

mAb 2C5-immunomicelles results in dramatically improved cancer cell killing by such 

micelles. Figure 20.3c presents the results of in vitro experiments with human breast cancer MCF-7 

cells that showed a clearly superior efficiency of paclitaxel-loaded 2C5-immunomicelles compared 

to paclitaxel-loaded plain or modified with a non-specific IgG micelles or free drug. 

In vivo experiments with LLC tumors-bearing mice revealed a dramatically enhanced tumor 

uptake of paclitaxel-loaded radiolabeled 2C5-immunomicelles compared to non-targeted micelles 

(Table 20.1). For tumor accumulation experiments in vivo, LLC tumors have been grown in mice 

subcutaneously injected with 20,000 LLC cells in 50 mL of 10 mM HBS into the left rear flank. 

When tumor diameters reached 3–7 mm (8–12 days post inoculation), the mice were injected with 

111 In-labeled plain or mAb 2C5-bearing paclitaxel-loaded PEG–PE micelles via the tail vein. At 

30 min or 2 h post injection, mice were sacrificed; tumors and muscle samples were collected and 

analyzed for the presence of 111 In radioactivity to estimate micelle accumulation in the tumor. To 

estimate the actual tumor accumulation of the drug, samples of tumor tissues from mice receiving 

ca. 100 mg paclitaxel per animal in plain PEG–PE micelles, in 2C5-immunomicelles, and the same 


O 2 N 

O 

C O 

O 2 N 

O 

C O 

O 

O 

C 

NO 2 

O 

O C 

Ligand 

NO2 HO 

N2H FIGURE 20.2 Schematic pattern of antibody (or other amino-containing ligand) attachment to activated PEG–PE micelles via the pNP group. 


OH 

414 



BT-20 

LLC 

EL4 

BT-20 

LLC 

EL4 

(a) (b) (c) 

1 2 3 

quantity of paclitaxel in Cremophor EL/ethanol/saline mixture (free paclitaxel) were analyzed 

30 min and 2 h post injection by HPLC. The results presented in the data in the Table 20.1 

clearly demonstrate that paclitaxel-loaded mAb 2C5-immunomicelles accumulate in tumors significantly 

better than drug-loaded plain micelles or free drugs and bring into tumors more drug 

(paclitaxel) than any other preparation. An enhanced accumulation of 2C5-targeted micelles over 

plain micelles in the tumor (up to 30%) was observed both at 30 min and at 2 h post injection 

indicating the specific recognition and tumor binding of 2C5-targeted immunomicelles. These data 

suggest the possibility that drug-loaded immunomicelles may also be better internalized by tumor 

cells similar to antibody-targeted liposome, 80 delivering more drug inside tumor cells that might be 

achieved in case of simple EPR effect-mediated tumor accumulation. By analyzing the absolute 

quantity of tumor-accumulated paclitaxel (HPLC 81 ) delivered by different drug formulations, it was 

shown that 2C5-immunomicelle was bringing into tumors substantially higher quantities of paclitaxel 

than in the case of paclitaxel-loaded non-targeted micelles or free drug formulation 

(Taxol w )(see Table 20.1). The difference in tumor accumulation of paclitaxel in immunomicelles 

compared to other drug preparations was larger at 2 h post injection than at 30 min post injection, 

explained by the accumulation of different preparations in different tumor compartments. Nontargeted 

paclitaxel formulations accumulate in the interstitial space of the tumor via the EPR effect 

Viable cells (%) 

100 

90 

80 

70 

60 

50 

40 

30 

20 

10 

0 

free paclitaxel 

plain micelles 

2C5-immunomicelles 

FIGURE 20.3 Specific properties of PEG–PE-based 2C5-immunomicelles in vitro. (a) Enhanced binding of 

rhodamine-labeled 2C5-immunomicelles with various cancer cells. (b) Enhanced binding of rhodaminelabeled 

drug(paclitaxel)-loaded 2C5-immunomicelles with various cancer cells. (c) Increased cytotoxicity 

of paclitaxel-loaded 2C5-immunomicelles toward cancer cells (with MCF7 cells as an example). 

TABLE 20.1 

Paclitaxel-Loaded PEG–PE-Based Micelles In Vivo 

Paclitaxel in Pain Micelles Paclitaxel in 2C5-Micelles Free Paclitaxel 

Tumor accumulation of 

micelles (% dose/g of tumor) 

4.15G0.36 6G0.10 n/a 

Tumor accumulation of 

paclitaxel (ng/g of tumor) 

200G30 850G40 130G25 

Inhibition of tumor growth (%) 35G5 74G34 23G4 

Accumulation measured 2 h post injection. Tumor growth inhibition measured over 10 days. Immunomicelles show the 

highest accumulation of all preparations, bring the maximum quantity of the drug into tumor, and allow for maximum 

inhibition of tumor growth. 


416 

and eventually drug could be cleared from the site after gradual micelle degradation. Contrary to 

that, paclitaxel-loaded 2C5-immunomicelles are internalized by cancer cells and keep the drug 

inside the tumor in a way similar to what was observed with drug-loaded anti-her2 immunoliposomes. 

80 The internalization by tumor cells may be highly therapeutically significant for many antitumor 

agents. For example, a much higher tumor regression was observed with a carrier capable of 

intracellular drug delivery for an equal doxorubicin dose delivered to the tumor. 80 It results in a 

higher therapeutic efficiency of paclitaxel-loaded mAb 2C5-micelles. The average weight of 

excised tumors in the group treated with paclitaxel in mAb 2C5-immunomicelles was approx. 

0.7 g compared to 1.6 g and 1.4 g in groups treated with the same quantity of paclitaxel as free 

drug or in plain PEG–PE micelles, respectively, (P!.05 in both cases). The weight of untreated 

tumors was around 2.0 g (Table 20.1). 

Recently, a poorly soluble photodynamic therapy (PDT) agent, meso-tetratphenylporphine 

(TPP), was effectively solubilized using non-targeted and tumor-targeted polymeric micelles 

prepared with PEG–PE. 82 Encapsulation of TPP into PEG–PE-based micelles and 2C5-immunomicelles 

(bearing an anti-cancer monoclonal 2C5 antibody) resulted in significantly improved anticancer 

effects of the drug at PDT conditions against murine (LLC, B16) and human (MCF-7, BT20) 

cancer cells in vitro. For this purpose, the cells were incubated for 6 h or 18 h with the TPP or TPPloaded 

PEG–PE micelles/immunomicelles and then light-irradiated for 30 min. The phototoxic 

effect depended on the TPP concentration and the light dose (the duration of light-irradiation) 

(see some data in Figure 20.4). An increased level of apoptosis was shown in the PDT-treated 

cultures. The attachment of the anti-cancer 2C5 antibodies to TPP-loaded micelles provided the 

maximum level of cell killing at a given time. Therefore, drug-loaded immunomicelles may also 

represent a useful formulation of the photosensitizer for practical PDT. 

Certain polymeric micelles possess a very high stability and ability to carry a variety of poorly 

soluble pharmaceuticals, and they can be used as targeted drug delivery systems into tumors. The 

tumor targeting can be achieved via the EPR effect by making micelles of stimuli-responsive 

amphiphilic block-copolymers or by attaching specific targeting ligand molecules (such as folate 

of transferrin) to the micelle surface. Tumor-specific immunomicelles can be prepared by coupling 

monoclonal antibody (such as antinucleosomal monoclonal antibody) molecules to pNP groups on 

the water-exposed terimini of the micelle corona-forming blocks. The micelle-coupled antibodies 

preserve their specific activity, and immunomicelles prepared with the use of the cancer-specific 

mAb 2C5 specifically bind to different cancer cells in vitro, demonstrating increased accumulation 

% survival 

100 

80 

60 

40 

20 

0 

0 

10 20 30 

TPP concentration (μg/ml) 


40 50 

FIGURE 20.4 Improved killing of cancer cells with the drug loaded into tumor-specific 2C5-immunomicelles 

under the PDT conditions. Phototoxicity of free TPP and TPP encapsulated in PEG–PE micelles and 2C5- 

PEG–PE-immunomicelles after light irradiation towards BT20 (human breast adenocarcinoma) cells; all 

preparations were pre-incubated with cells for 18 h): Cfree TPP; $TPP in PEG–PE micelles; %TPP in 

2C5-PEG–PE-immunomicelles. 



in experimental tumors in vivo. Being loaded with poorly soluble anti-cancer drugs (such as 

paclitaxel, camptothecin, 83 or photodynamic therapy agents), 82 mAb 2C5-immunomicelles demonstrate 

significantly increased cytotoxicity toward tumor cells in vitro and in vivo. 

REFERENCES 

1. Müller, R. H., Colloidal Carriers for Controlled Drug Delivery and Targeting, Wissenschaftliche 

Verlagsgesellschaft/CRC Press, Stuttgart, Germany/Boca Raton, 1991. 

2. Cohen, 2. S. and Bernstein, H., Eds., Microparticulate Systems for the Delivery of Proteins and 

Vaccines, Marcel Dekker, Inc., New York, 1996. 

3. Lasic, D. D. and Martin, F., Eds., Stealth Liposomes, CRC Press, Boca Raton, 1995. 

4. Torchilin, V. P. and Trubetskoy, V. S., Which polymers can make nanoparticulate drug carriers longcirculating?, 


5. Palmer, T. N. et al., The mechanism of liposome accumulation in infarction, Biochim. Biophys. Acta, 

797, 363, 1984. 


review, J. Control. Rel., 65, 271, 2000. 

7. Torchilin, V. P., Polymer-coated long-circulating microparticular pharmaceuticals, J. Microencapsul., 

15, 1, 1998. 

8. Lipinski, C. A. et al., Experimental and computational approaches to estimate solubility and permeability 

in drug discovery and development settings, Adv. Drug Deliv. Rev., 46, 3, 2001. 

9. Fernandez, A. M. et al., N-Succinyl-(beta-alanyl-L-leucyl-L-alanyl-L-leucyl)doxorubicin: An extracellularly 

tumor-activated prodrug devoid of intravenous acute toxicity, J. Med. Chem., 44, 3750, 

2001. 

10. Yalkowsky, S. H., Ed,. Techniques of Solubilization of Drugs, Marcel Dekker, New York, 1981. 

11. Yokogawa, K. et al., Relationships in the structure-tissue distribution of basic drugs in the rabbit, 

Pharm. Res., 7, 691, 1990. 

12. Hageluken, A. et al., Lipophilic beta-adrenoceptor antagonists and local anesthetics are effective 

direct activators of G-proteins, Biochem. Pharmacol., 47, 1789, 1994. 

13. Shabner, B. A. and Collings, J. M., Eds., Cancer Chemotherapy: Principles and Practice, J.B. 

Lippincott Co., Philadelphia, 1990. 

14. Lasic, D. D. and Papahadjopoulos, D., Medical Applications of Liposomes, Elsevier, New York, 1998. 

15. Thompson, D. and Chaubal, M. V., Cyclodextrins (CDS)—excipients by definition, drug delivery 

systems by function (part I: Injectable applications), Drug Deliv. Technol., 2, 34, 2000. 

16. Mittal, K. L. and Lindman, B., Eds., Surfactants in Solution, Vol. 1–3, Plenum Press, New York, 1991. 

17. Lasic, D. D., Mixed micelles in drug delivery, Nature, 355, 279, 1992. 

18. Kwon, G. S. and Kataoka, K., Block copolymer micelles as long-circulating drug vehicles, Adv. Drug 

Deliv. Rev., 16, 295, 1995. 

19. Kwon, G. S., Diblock copolymer nanoparticles for drug delivery, Crit. Rev. Ther. Drug Carrier Syst., 

15, 481, 1998. 

20. Kabanov, A. V., Batrakova, E. V., and Alakhov, V. Y., Pluronic block copolymers as novel polymer 

therapeutics for drug and gene delivery, J. Control. Rel., 82, 189, 2002. 

21. Jones, M. and Leroux, J., Polymeric micelles-a new generation of colloidal drug carriers, Eur. 

J. Pharm. Biopharm., 48, 101, 1999. 

22. Torchilin, V. P., Structure and design of polymeric surfactant-based drug delivery systems, J. Control. 

Rel., 73, 137, 2001. 

23. Zhang, L. and Eisenberg, A., Multiple morphologies of “crew-cut” aggregates of polystyrene-bpoly(acrylic 

acid) block copolymers, Science, 268, 1728, 1995. 

24. Gref, R. et al., Biodegradable long-circulating polymeric nanospheres, Science, 263, 1600, 1994. 

25. Gref, R. et al., The controlled intravenous delivery of drugs using PEG-coated sterically stabilized 

nanospheres, Adv. Drug Deliv. Rev., 16, 215, 1995. 

26. Kabanov, A. V. et al., The neuroleptic activity of haloperidol increases after its solubilization in 

surfactant micelles, FEBS Lett., 258, 343, 1989. 


418 


27. Hagan, S. A. et al., Polylactide-poly(ethelene glycol) copolymers as drug delivery systems 1. Caracterization 

of water dispersible micelle-forming systems, Langmuir, 12, 2153, 1996. 

28. Inoue, T. et al., An AB block copolymers of oligo(methyl methacrylate) and poly(acrylic acid) for 

micellar delivery of hydrophobic drugs, J. Control. Rel., 51, 221, 1998. 

29. Hunter, R. J., Foundations of Colloid Science, Vol. 1, Oxford University Press, New York, 1991. 

30. Kuntz, R. M. and Saltzman, W. M., Polymeric controlled delivery for immunization, Trends Biotech., 

15, 364, 1997. 

31. Torchilin, V. P. et al., Amphiphilic vinyl polymers effectively prolong liposome circulation time 

in vivo, Biochim. Biophys. Acta, 1195, 181, 1994. 

32. Torchilin, V. P. et al., New sunthetic amphiphilic polymers for steric protection of liposomes in vivo, 

J. Pharm. Sci., 84, 1049, 1995. 

33. Miller, D. W. et al., Interactions of pluronic block copolymers with brain microvessel endothelial 

cells: Evidence of two potential pathways for drug absorption, Bioconjug. Chem., 8, 649, 1997. 


supramolecular assembly with poly(ethylene glycol)–poly(L-lysine) block copolymer, J. Pharm. Sci., 

87, 160, 1998. 

35. Trubetskoy, V. S. et al., Block copolymer of polyethylene glycol and polylysine as a carrier of organic 

iodine: Design of a long circulating particulate contrast medium for x-ray computed tomography, 

J. Drug Target., 4, 381, 1997. 

36. Yokoyama, M. et al., Characterization and anticancer activity of the micelle-forming polymeric 

anticancer drug adriamycin-conjugated poly(ethylene glycol)–poly(aspartic acid) block copolymer, 

Cancer Res., 50, 1693, 1990. 

37. Harada, A. and Kataoka, K., Novel polyion complex micelles entrapping enzyme molecules in the 

core. Preparation of narrowly-distributed micelles from lysozyme and poly(ethylene glycol)–poly 

(aspartic acid) block copolymer in aqueous medium, Macromolecules, 31, 288, 1998. 


polymeric drug indomethacin-incorporated poly(ethylene oxide)–poly(-benzyl L-aspartat) block 

copolimer micelles, J. Pharm. Sci., 85, 85, 1996. 

39. Kwon, G. S. et al., Block copolymer micelles for drug delivery: Loading and release of doxorubicin, 

J. Control. Rel., 48, 195, 1997. 

40. Kim, S. Y. et al., Metoxy poly(ethylene glucol) and 3-caprolactone amphiphilic block copolymeric 

micelle containing indomethacin, II. Micelle formation and drug release behaviors, J. Control. Rel., 

51, 13, 1998. 

41. Allen, C. et al., Polycaprolactone-b-poly(ethylene oxide) block copolymer micelles as a novel drug 

delivery vehicle for neurotrophic agents FK506 and L-685,818, Bioconjug. Chem., 9, 564, 1998. 

42. Ramaswamy, M. et al., Human plasma distribution of free paclitaxel and paclitaxel associated with 

diblock copolymers, J. Pharm. Sci., 86, 460, 1997. 

43. Trubetskoy, V. S. and Torchilin, T. P., Use of polyoxyethylene–lipid conjugates as long-circulating 

carriers for delivery of therapeutic and diagnostic agents, Adv. Drug Deliv. Rev., 16, 311, 1995. 

44. Klibanov, A. L. et al., Amphipathic polyethyleneglycols effectively prolong the circulation time of 

liposomes, FEBS Lett., 268, 235, 1990. 

45. Lasic, D. D. et al., Valentincic, phase behavior of “stealth–lipid” decithin mixtures, Period. Biol., 93, 

287, 1991. 

46. Ray, R. et al., Handbook of Pharmaceutical Excipients, APhA Publications, Washington, 2003. 

47. Trubetskoy, V. S. and Torchilin, V. P., Polyethylene glycol based micelles as carriers of therapeutic 

and diagnostic agents, S.T.P. Pharma Sci., 6, 79, 1996. 

48. Gao, Z. et al., Diacyl-polymer micelles as nanocarriers for poorly soluble anticancer drugs, Nano 

Lett., 2, 979, 2002. 

49. Wang, J. et al., Preparation and in vitro synergistic anticancer effect of Vitamin K# and 1,8-diazabicyclo[5,4,0]undec-7-ene 

in poly(ethylene glycol)-diacyllipid micelles, Int. J. Pharm., 272, 129, 2004. 


delivery, Coll. Surf. B: Biointerf., 16, 1, 1999. 

51. Alakhov, V. Yu. and Kabanov, A. V., Block copolymeric biotransport carriers as versatile vehicles for 

drug delivery, Expert. Opin. Invest. Drugs, 7, 1453, 1998. 



52. Matsumura, Y. et al., Reduction of the side effects of an antitumor agent, KRN5500, by incorporation 

of the drug into polymeric micelles, Jpn. J. Cancer Res., 90, 122, 1999. 

53. Wang, J., Mongayt, D., and Torchilin, V. P., Polymeric micelles for delivery of poorly soluble drugs: 

Preparation and anticancer activity in vitro of paclitaxel incorporated into mixed micelles based on 

poly(ethylene glycol)–lipid conjugate and positively charged lipids, J. Drug Target., 13, 73, 2005. 

54. Weissig, V., Whiteman, K. R., and Torchilin, V. P., Accumulation of liposomal- and micellar-bound 

protein in solid tumor, Pharm. Res., 15, 1552, 1998. 

55. Lukyanov, A. N. et al., Polyethylene glycol-diacyllipid micelles demonstrate increased accumulation 

in subcutaneous tumors in mice, Pharm. Res., 19, 1424, 2002. 


cutoff size, Cancer Res., 55, 3752, 1995. 

57. Hobbs, S. K. et al., Regulation of transport pathways in tumor vessels: Role of tumor type and 

microenvironment, Proc. Natl Acad. Sci. USA, 95, 4607, 1998. 

58. Monsky, W. L. et al., Augmentation of transvascular transport of macromolecules and nanoparticles 

in tumors using vascular endothelial growth factor, Cancer Res., 59, 4129, 1999. 

59. Helmlinger, G. et al., Interstitial pH and pO2 gradients in solid tumors in vivo: High-resolution 

measurements reveal a lack of correlation, Nat. Med., 3, 177, 1997. 

60. Tannock, I. T. and Rotin, D., Acid pH in tumors and its potential for therapeutic exploitation, Cancer 

Res., 49, 4373, 1989. 

61. Kwon, G. S. and Okano, T., Soluble self-assembled block copolymers for drug delivery, Pharm. Res., 

16, 597, 1999. 

62. Cammas, S. et al., Thermorespensive polymer nanoparticles with a core-shell micelle structure as site 

specific drug carriers, J. Control. Rel., 48, 157, 1997. 

63. Chung, J. E. et al., Effect of molecular architecture of hydrophobically modified poly(N-isopropylacrylamide) 

on the formation of thermoresponsive core-shell micellar drug carriers, J. Control. Rel., 

53, 119, 1998. 

64. Kohori, F. et al., Preparation and characterization of thermally responsive block copolymer micelles 

comprising poly(N-isopropylacrylamide-b-DL-lactide), J. Control. Rel., 55, 87, 1998. 

65. Meyer, O., Papahadjopoulos, D., and Leroux, J. C., Copolymers of N-isopropylacrylamide can trigger 

pH sensitivity to stable liposomes, FEBS Lett., 41, 61, 1998. 

66. Le Garrec, D. et al., Optimizing pH-responsive polymeric micelles for drug delivery in a cancer 

photodynamic therapy model, J Drug Target., 10, 429, 2002. 

67. Yoo, H. S., Lee, E. A., and Park, T. G., Doxorubicin-conjugated biodegradable polymeric micelles 

having acid-cleavable linkages, J Control. Rel., 82, 17, 2002. 

68. Bae, Y. et al., Multifunctional polymeric micelles with folate-mediated cancer cell targeting and pHtriggered 

drug releasing properties for active intracellular drug delivery, Mol. Biosyst., 1, 242–250, 

2005. 

69. Yokoyama, M. et al., Stabilization of disulfide linkage in drug-polymer-immunoglobulin conjugate by 

microenvironmental control, Biochem. Biophys. Res. Commun., 164, 1234–1239, 1989. 

70. Nagasaki, Y. et al., Sugar-installed block copolymer micelles: Their preparation and specific 

interaction with lectin molecules, Biomacromolecules, 2, 1067, 2001. 

71. Vinogradov, S. et al., Polyion complex micelles with protein-modified corona for receptor-mediated 

delivery of oligonucleotides into cells, Bioconjug. Chem., 10, 851, 1999. 

72. Leamon, C. P., Weigl, D., and Hendren, R. W., Folate copolymer-mediated transfection of cultured 

cells, Bioconjug. Chem., 10, 947, 1999. 

73. Leamon, C. P. and Low, P. S., Folate-mediated targeting: From diagnostics to drug and gene delivery, 

Drug Discov. Today, 6, 44, 2001. 

74. Lee, E. S., Na, K., and Bae, Y. H., Polymeric micelle for tumor pH and folate-mediated targeting, J 

Control. Rel., 91, 103, 2003. 

75. Torchilin, V. P. et al., Immunomicelles: Targeted pharmaceutical carriers for poorly soluble drugs, 

Proc. Natl Acad. Sci. USA, 100, 6039, 2003. 

76. Torchilin, V. P. et al., p-Nitrophenylcarbonyl–PEG–PE–liposomes: Fast and simple attachment of 

specific ligands, including monoclonal antibodies, to distal ends of PEG chains via p-nitrophenylcarbonyl 

groups, Biochim. Biophys. Acta, 1511, 397, 2001. 


420 


77. Gao, Z. et al., PEG–PE/phosphatidylcholine mixed immunomicelles specifically deliver encapsulated 

taxol to tumor cells of different origin and promote their efficient killing, J. Drug Target., 11, 87, 2003. 

78. Iakoubov, L. Z. and Torchilin, V. P., A novel class of antitumor antibodies: Nucleosome-restricted 

antinuclear autoantibodies (ANA) from healthy aged nonautoimmune mice, Oncol. Res., 9, 439, 1997. 

79. Torchilin, V. P., Iakoubov, L. Z., and Estrov, Z., Therapeutic potential of antinuclear autoantibodies in 

cancer, Cancer Ther., 1, 179, 2003. 

80. Park, J. W. et al., Tumor targeting using anti-her2 immunoliposomes, J. Control. Rel., 74, 95, 2001. 

81. Sharma, A., Conway, W. D., and Straubinger, R. M., Reversed-phase high-performance liquid chromatographic 

determination of taxol in mouse plasma, J. Chromatogr. B. Biomed. Appl., 655, 315, 

1994. 

82. Roby, A., Erdogan, S., and Torchilin, V. P., Solubilization of poorly soluble PDT agent, mesatetraphenylporphin, 

in plain or immunotargeted PEG–PE micelles results in dramatically improved 

cancer cell killing in vitro. Eur. J. Pharm. Biopharm., 62(3):235–240,2006. 

83. Mu, L., Elbayoumi, T. A., and Torchilin, V. P., Mixed micelles made of poly(ethylene glycol)– 

phosphatidylethanolamine conjugate and d-alpha-tocopheryl polyethylene glycol 1000 succinate as 

pharmaceutical nanocarriers for camptothecin, Int. J. Pharm., 306(1–2): 142–149, 2005. 


21 

CONTENTS 

Combined Cancer Therapy by 

Micellar-Encapsulated Drugs 

and Ultrasound 


21.1 Introduction ....................................................................................................................... 421 

21.2 Polymeric Micelles as Drug Carriers................................................................................ 422 

21.2.1 Advantages and Shortcomings of Polymeric Micelles as Drug Carriers .......... 422 

21.2.2 Polymeric Micelles in Clinical Trials ................................................................. 423 

21.3 Ultrasound in Drug Delivery ............................................................................................ 425 

21.3.1 Introduction to Ultrasound in Medicine ............................................................. 425 

21.3.2 Biological Effects of Ultrasound......................................................................... 425 

21.3.2.1 Cellular Level ..................................................................................... 425 

21.3.2.2 Systemic Level ................................................................................... 427 

21.3.3 Ultrasound-Induced Drug Release from Polymeric Micelles ............................ 428 

21.3.4 Ultrasound-Enhanced Intracellular Drug Uptake ............................................... 428 

21.4 In Vivo Evaluation of the Drug/Micelle/Ultrasound Tumor-Targeting Modality........... 430 

21.4.1 Ultrasound-Induced Drug Release from Micelles In Vivo ................................ 430 

21.4.2 Biodistribution of Polymeric Micelles and Micellar-Encapsulated Drugs ........ 430 

21.4.3 Effect of the Time of Ultrasound Application.................................................... 432 

21.4.4 Inhibition of Tumor Growth Using Micelle/Ultrasound Drug Delivery ........... 432 

21.4.4.1 Drug-Sensitive Tumors ...................................................................... 432 

21.4.4.2 Drug Resistant Tumors ...................................................................... 433 

21.4.4.3 Poorly Vascularized Tumors.............................................................. 435 

21.4.5 Mechanisms of Micelle/Ultrasound Bioeffects .................................................. 436 

21.5 Combining Ultrasound Imaging and Treatment ............................................................... 437 

21.5.1 Dual-Modality Contrast Agents/Drug Carrier Systems...................................... 437 

21.6 Conclusions ....................................................................................................................... 438 

References..................................................................................................................................... 438 


Chemotherapy is often complicated by toxic side effects of anti-cancer drugs. Also, effective 

chemotherapy regimens are frequently hindered by the resistance of tumor cells to one or more 

drugs [cross-resistance or multidrug resistance (MDR)]. The MDR is either inherent or acquired in 


421

422 

the process of chemotherapy. Developing new tumor-localized chemotherapeutic modalities for 

treatment of drug-sensitive and MDR tumors is a major goal of current research in academia and 

industry. Towards this end, nanoparticle-based drug delivery combined with a localized drug 

release from carrier in the tumor volume is a promising approach to targeted chemotherapy. 

This chapter reviews the state of the art in drug targeting to tumors using polymeric micelles as 

drug carriers and tumor-localized ultrasound as a means to trigger drug release from micelles in the 

tumor volume. Ultrasound enhances the intracellular drug uptake and sensitizes MDR tumors to the 

action of conventional drugs. 

21.2 POLYMERIC MICELLES AS DRUG CARRIERS 


The polymeric micelles are formed by hydrophobic–hydrophilic block copolymers at concentrations 

above the critical micelle concentration (CMC). Below the CMC, block copolymer 

molecules exist in a solution in the form of individual molecules (unimers). At concentrations 

above the CMC, copolymer molecules self-assemble into dense micelles with hydrophobic cores 

and hydrophilic corona. The amphiphilic character of block copolymer micelles, their size 

(approximately 10–150 nm), and surface properties provide for a high drug loading capacity and 

long circulation time in the vascular system. This makes them attractive drug carriers. 1,2 

21.2.1 ADVANTAGES AND SHORTCOMINGS OF POLYMERIC MICELLES AS DRUG CARRIERS 

Drug encapsulation in micelles decreases systemic concentration of drug and diminishes intracellular 

drug uptake by normal cells, thus reducing unwanted drug interactions with healthy tissues. 

The important advantage offered by polymeric micelles is a so-called “enhanced penetration and 

retention (EPR) effect” 3 that provides for a selective accumulation of micellar-encapsulated drugs 

in tumor interstitium; in turn, this offers prospects for controlled drug release in the tumor volume. 

The simplicity of micelle formation by self-assembly of amphiphilic block copolymer 

molecules and drug encapsulation by physical mixing rather than chemical conjugation are extremely 

attractive features of polymeric micelles. 

Another important advantage offered by polymeric micelles is related to solubilization of drugs 

with poor aqueous solubility. Most anti-cancer drugs are lipophilic; thus, they have low aqueous 

solubility; conventional solubilizing agents, currently in use for the formulation of low-solubility 

drugs, are usually very toxic. Use of polymeric micelles as solubilizing agents results in dramatically 

increased aqueous solubility, and substantially decreased systemic toxicity of clinical 

formulations of paclitaxel (PTX), doxorubicin (DOX), and many other anti-cancer drugs. 4–18 

Various polymeric micellar systems have been designed to optimize important parameters of 

drug performance, such as aqueous solubility, on-demand release, and biological distribution. In 

this context, the most thoroughly studied are the micelles that comprise of poly(ethylene oxide) 

(PEO) as a hydrophilic block (or blocks). Examples of these include PEO-b-poly(propylene oxide)b-PEO 

copolymers (Pluronic w ), PEO-b-polyesters, and PEO-b-poly(amino acid)s. 

The application of biodegradable, pH-sensitive micelles like those of poly(ethylene glycol)-copoly(L-lactide) 

(PEG–PLLA) or poly(ethylene glycol)-co-poly(caprolactone) (PEG–PCL), may be 

especially attractive. If internalized by tumor cells via endocytosis, the micelles end up in the acidic 

environment of endosomes and lysosomes, where their accelerated degradation will release the 

encapsulated drug. 

In vitro drug release from micellar systems is usually assessed either by dialysis through 

semipermeable membranes or by gel-permeation chromatography. Unfortunately, the information 

provided by these techniques is not directly pertinent to the in vivo behavior of drug-loaded 

polymeric micelles. The systemic injection of micellar formulations is associated with substantial 

dilutions and sink conditions. Because micelle formation is thermodynamically driven, when 

diluted below the CMC, micelles dissociate into individual molecules (unimers), thus releasing 


Combined Cancer Therapy by Micellar-Encapsulated Drugs and Ultrasound 423 

their drug load. To prevent premature micelle dissociation and drug release upon injection, micellar 

systems should have low CMC, glassy cores, or cross-linked cores. This may be achieved by 

induction of strong hydrophobic interactions or hydrogen bonds in the micelle cores. 19–21 

21.2.2 POLYMERIC MICELLES IN CLINICAL TRIALS 

Many micellar drug delivery systems showed promising results in in vitro and animal experiments, 

but showed disappointing behavior in clinical trials. 

The first clinical trials of micellar drug delivery systems began in this decade. Physically 

encapsulated DOX in PEO5000-b-P(Asp)4000-DOX micelles (NK911 formulation) entered clinical 

trials in Japan in 2001. 22 In animal studies, NK911 demonstrated about 30-fold higher area-underthe-curve 

(AUC) in plasma, higher tumor drug levels, lower toxicity, and higher therapeutic 

activity. However, in clinical trials, NK911 showed the same spectrum of side effects as free 

DOX, despite somewhat more favorable pharmacokinetic parameters. The discrepancy between 

the expected and observed clinical results was most likely caused by a premature drug release from 

micelles due to a very strong dilution of infused micellar solutions in the human circulatory system 

(volume: w5 L) when compared to that of mice (volume: w2 mL). A drop-like infusion of micellar 

drug delivery systems appears to have no prospect due to the dilution problem. Only bolus injections 

of reasonably stable micellar systems could provide for micelle preservation and drug 

retention in circulation. 

Aqueous Pluronic w solutions are another polymeric system suggested for DOX delivery, 

especially for drug resistant cancers. Pluronic w copolymers in unimeric concentrations were 

shown to hyper-sensitize drug resistance carcinoma cells and enhance response of resistant 

tumors to chemotherapeutic agents. 23–31 Thereasonbehindthisphenomenonisrelatedtoan 

increased drug uptake caused by deactivation of drug efflux pumps; the latter process is associated 

with Pluronic w -induced energy depletion in resistant tumor cells. 32 In addition, favorable changes 

in the intracellular trafficking of DOX occur in the MDR cells under Pluronic w action. 30,33,34 

Pluronic w formulation of DOX, known as SP1049C, was thoroughly investigated in animal 

studies by Alakhov et al. 35 Although this formulation is probably partly micellar, it is not expected 

to retain micellar structure upon intravenous injections. In vivo, DOX and Pluronic w unimers 

mostlikely move along independent routes. Not unexpectedly, the toxic level of DOX in 

SP1049C formulation was shown to be similar to that of free DOX. The results of pharmacokinetic 

and biodistribution experiments on SP1049C in Lewis lung carcinoma bearing C57BL/6 mice 

showed a very moderate increase in plasma AUC. In anti-tumor efficacy studies using hematological 

and solid animal tumor models, SP1049C was shown to be more effective than free DOX at 

an equal dose of 5 mg/kg. 30 

SP1049C entered preclinical and clinical trials in Canada in 1999. 36 The toxicity profiles of free 

DOX and Pluronic w formulation were identical, except for the histopathological changes in skin 

and thymus, that were less pronounced with SP1049C. In 2004, Danson et al. reported on the results 

of phase-I clinical trials of SP1049C. 36 In this study, SP1049C was given to patients with histologically 

proven cancer refractory to conventional treatments. The dose-limiting toxicity (DLT) in 

humans, i.e., myelosuppression, was reached at 90 mg/m 2 of SP1049C; therefore, a dose of 70 mg/ 

m 2 was recommended for future trials. The pharmacokinetic profile of SP1049C was similar to that 

of a clinical DOX formulation, except for a slower terminal clearance stage. SP1049C induced 

temporal partial response in several patients with advanced solid tumors. The results were 

considered promising, and the phase-II trials were initiated; their first results were recently reported 

by Valle et al. 37 In phase-II trials, SP1049C was evaluated in patients with less severe forms of 

cancer. Patients received 75 mg/m 2 of SP1049C (30 min intravenous infusion), four times a week, 

for up to six cycles. The authors reported partial responses in some patients after 4–6 treatment 

cycles. However, data showed appearance of hematological and nonhematological (especially 


424 


cardiac) toxicities, presumably because DOX was not encapsulated and, therefore, not prevented 

from the attack on the heart muscle. 

PTX is another potential drug candidate for the delivery in polymeric micelles. PTX is a highly 

hydrophobic molecule; for clinical use, PTX is solubilized with the aid of a significant amount of 

Cremophor EL and ethyl alcohol. 38 These formulations (Taxol w or PTX Injections) cause significant 

systemic toxicity. Considerable efforts of various groups have been dedicated to developing 

micellar formulations of PTX. 4–6,9,39 However, except for the Japanese formulation NK105 and 

Genexol w (see below), no other formulations have progressed to clinical trials due to a rapid loss of 

drug from micellar carriers, resulting in the absence of advantages over clinical formulations. 

In 2001, Kim et al. reported on PTX-loaded PEO-b-poly(D,L-lactide) (PEO2000-b-PDLLA1750) 

micelles (Genexol w -PM) prepared by a solvent evaporation technique. 39 The maximum tolerated 

doses (MTDs) of Genexol w -PM and Taxol w for the intravenous (i.v.) administration in nude mice 

were 60 and 20 mg/kg, respectively. This allowed a threefold higher administration dose of drug. In 

animal studies, Genexol w -PM showed a better anti-tumor effect in ovarian and breast cancer mice 

models than clinical Taxol w . 

However, despite an increase in the administered dose, plasma AUC was 20-fold lower for the 

micellar formulation, which was attributed to a fast release of drug from the micelles. Another 

factor involved in the fast elimination of micellar PTX from the circulation could be related to the 

uptake by the cells of the reticulo-endothelial system (RES). In all organs except plasma, PTX 

biodistribution was comparable for a clinical formulation or Genexol-PM. In phase-I human trials, 

micellar-encapsulated PTX allowed administration of a higher drug dose, but resulted in a lower 

plasma AUC and a shorter PTX half life. 40 

Another type of polymeric micelles suggested for PTX delivery comprise PDLLA as a coreforming 

block and poly(N-vinylpyrrolidone) (PVP) as a shell-forming block. 41 The in vivo evaluation 

of this system showed results similar to those for Genexol w -PM in regards to a reduced 

plasma AUC. 42 For the micellar formulation, the AUC was also reduced in liver, kidney, spleen, 

and heart, but was not significantly changed in the tumor. For the polymeric micellar formulation, 

MTD was not reached even at 100 mg/kg, whereas the MTD of Taxol w was 20 mg/kg. This allowed 

a threefold increase of the therapeutic dose of PTX, resulting in a significant increase of the 

anti-tumor activity of the polymeric system compared to Taxol w . 

Polymeric micellar PTX formulation NK105 reported by Hamaguchi et al. 43 comprises 

micelle-forming block copolymer poly(ethelene oxide)-b-poly(4-phenyl-1-butanoate)-L-apartamide 

(PEO-b-PPBA). This formulation demonstrated stealth and targeting properties. NK105 

manifested a 86-fold increase in PTX plasma AUC compared to Taxol w , accompanied by a 

25-fold increase for drug AUC in tumor, and a higher anti-tumor activity in C-26 tumor-bearing 

mice. At a PTX dose of 100 mg/kg, tumors were completely resolved after a single administration 

of NK105. This formulation is currently in phase-I clinical trials in Japan. 

In summary, only a few polymeric micellar formulations have demonstrated success in clinical 

trials. Polymeric micelles as drug carriers are either too stable, therefore incapable of providing 

adequate drug release at the tumor site, or too unstable, thus prematurely releasing their 

drug payload. 

To overcome these complications, we have developed Pluronic w -based micellar systems that 

withstand the destabilizing effects of the biological environment and effectively target tumors while 

retaining their drug load. 44 Various routes of micelle stabilization were tested for Pluronic 

micelles; 19 the optimal stability, low toxicity, and long circulation time of micelle-encapsulated 

DOX was achieved by forming mixed micelles of Pluronic w and PEO-diacyl phospholipids 44 or by 

introducing a small concentration of oil in the micelle core to increase core hydrophobicity. 19 The 

on-demand release of DOX from micelles and an effective intracellular drug uptake by the tumor 

cells was achieved by local ultrasound treatment of the tumor at a particular time after the drug 

injection. 44 



Important advantages of ultrasound are that it is noninvasive, can penetrate deep into the 

interior of the body, and can be focused and carefully controlled. Therapeutic tumor treatment 

by ultrasound requires tumor imaging by diagnostic ultrasound prior to therapy. During the last 

decade, combining diagnostic and therapeutic ultrasound in the same instrument has attracted ever 

growing attention. Developing dual-modality systems that combine properties of ultrasound 

contrast agents and drug carriers is a timely task. 

21.3 ULTRASOUND IN DRUG DELIVERY 

21.3.1 INTRODUCTION TO ULTRASOUND IN MEDICINE 

Ultrasound consists of pressure waves generated by piezoelectric transducers, with frequencies at or 

greater than 20 kHz. Like optical or audio waves, ultrasonic waves can be focused, reflected, 

refracted, and propagated through a medium, thus allowing the waves to be directed to and 

focused on a particular volume of tissue in a spatially and temporally controlled manner. Ultrasound 

technology allows for a high degree of spatial and dynamic control due to a favorable range 

of energy penetration characteristics in soft tissue and the ability to shape energy 

deposition patterns. 

In medicine, ultrasound is used as either a diagnostic or a therapeutic modality. The main 

advantage of ultrasound is its noninvasive nature; the transducer is placed in contact with a waterbased 

gel or water layer on the skin, and no insertion or surgery is required. 

In current therapeutic applications, ultrasound is used mostly for tissue thermoablation. 

Progress in imaging techniques has allowed ablating large tumor masses by high intensity 

focused ultrasound (HIFU). HIFU instruments noninvasively deliver high intensity ultrasound 

energy to predetermined volumes of the body. 

However, the ablative techniques require delivering high acoustic energy to a well delineated 

volume in the body over a substantial time period (up to several hours), which is associated with a 

potential risk of ablating surrounding tissues due to a patient’s breathing or motion. Motion 

correction is a serious concern in HIFU applications. In addition, some locations in the body are 

screened by bones and not accessible for HIFU treatment. 

In our lab, we have been developing a different mode of application of therapeutic ultrasound. 

In this modality, tumor cell killing is based on the ultrasound-enhanced cytotoxic action of tumortargeted 

chemotherapeutic drugs rather than the mechanical or thermal action of ultrasound. 45–56 

This technology requires order of magnitude lower ultrasound energy, reducing the risk of coagulative 

necrosis of healthy tissues associated with HIFU. 

21.3.2 BIOLOGICAL EFFECTS OF ULTRASOUND 

21.3.2.1 Cellular Level 

The effect of ultrasound on the cellular level is usually attributed to formation of plasma membrane 

defects (sonoporation) induced by cavitating microbubbles; the formation and collapse of microbubbles 

is called inertial cavitation. This process has been investigated using a voltage clamp 

technique. 57–59 It was shown that sonoporation was substantially enhanced by ultrasound contrast 

agents that nucleate inertial cavitation. 60 The degree of membrane damage increased with 

increasing duty cycle of pulsed ultrasound. 59 Some membrane defects are resealed when ultrasound 

is turned off (this is called transient sonoporation); a key role in membrane resealing is played by 

Ca 2C . 58 

In our experiments, ultrasound bioeffects on the cellular level were compared for the drugsensitive 

and multidrug resistant (MDR) A2780 ovarian carcinoma cells in the presence or absence 

of DOX. 61 Focused 1.1-MHz ultrasound was applied at varying peak negative pressures and DOX 

concentrations. Cell viability, a degree of membrane damage, DOX uptake, and cell proliferation 


426 

were measured. Cell incubation with the nucleic acid stain Sytox allowed discriminating between 

the cells experiencing various degrees of plasma membrane damage. Fluorescence of Sytox is 

negligible in the extracellular environment; when Sytox enters the cells and intercalates the 

DNA, its fluorescence increases by several orders of magnitude. For the present application it is 

important that Sytox does not penetrate into the cells with intact membranes; however, when cell 

membranes are damaged, Sytox enters the cells and intercalates the DNA, which results in a 

dramatic fluorescence increase. 

Major results of our experiments on ultrasound bioeffects are listed below. 

† Under the action of ultrasound, cells with various degrees of membrane damage, intermediate 

between viable and dead cells, were generated. 

† At the same delivered ultrasound energy, higher ultrasound pressure caused more severe 

membrane damage. 

† A proliferation rate of sonicated multidrug resistant ovarian carcinoma A2780/AD cells 

with compromised membranes was significantly lower than the proliferation rate of the 

intact cells (Figure 21.1). 

† The MDR A2780/AD cells manifested a higher susceptibility to ultrasound than parental 

drug-sensitive A2780 cells (Figure 21.2). 

† Introduction of DOX in the incubation medium increased cell susceptibility to the 

ultrasound action. For the same ultrasound energy, the cells sonicated in the presence 

of DOX manifested significantly higher membrane damage than the cells without DOX 

(Figure 21.3). For the same initial DOX concentration in the incubation medium, the 

effect of DOX was stronger at 378C than at room temperature suggesting that an 

increased cell susceptibility to ultrasound was caused by DOX internalized by the 

cells. We hypothesize that the effect of DOX may be associated with a destabilizing 

action on the DNA molecule at the intercalation sites. 

Very interesting data on the effect of ultrasound on the gene expression in multidrug resistant 

Hep2D/ADR hepatic carcinoma cells were recently obtained in Chongqing Medical University by 

Shao and Wu (Wu, F., personal communication). These researchers observed a noticeable suppression 

of the MDR1 gene expression in the cells sonicated by 0.8-MHz ultrasound; the BCL2/BAX 

expression ratio was also substantially reduced, resulting in an enhanced apoptosis. The last effect 

was significantly stronger in the MDR compared to parental cells. 

Percent survival/growth 

70 

60 

50 

40 

30 

20 

10 

0 

Immediate viability 


no DOX 

[DOX]= 1 ug/ml 

[DOX]= 20 ug/ml 

Proliferation 

FIGURE 21.1 Immediate cell viability (left group of bars) and proliferation at 378C for 72 h (right group 

of bars) for the A2780/AD cells sonicated (0.55 MPa, 15 s) in the presence or absence of various concentrations 

of DOX. (Adapted from Kamaev, P., Christensen, D. A., and Rapoport, N., unpublished results.) 



Viability, % 

These clinically important findings can be used for effective therapy of multidrugresistant 

tumors. 

21.3.2.2 Systemic Level 

In living organisms, the bioeffects of ultrasound are not limited to sonoporation. On a systemic 

level, a very important role may be played by enhanced extravasation. Because of the adverse 

effects that this may have in diagnostic ultrasound, extensive work has been done on studying 

extravasation at various locations in the body and from various blood vessels, especially in 

the process of heart imaging. Extravasation was substantially enhanced by ultrasound contrast 

agents. Enhanced leakage of the die and erythrocytes from heart vessels was observed in several 

studies. 62–64 Erythrocyte extravasation under the action of contrast-enhanced ultrasound was also 

observed in the kidney; it was shown that the effect of pulsed ultrasound was stronger than that of 

continuous wave (CW) ultrasound. 65 

Although enhanced extravasation is an undesirable effect in ultrasound imaging, it may play a 

very positive role in drug delivery, especially for the local delivery of macromolecular drugs or 

Viabilty, % 

100 

90 

80 

70 

60 

50 

40 

30 

20 

10 

0 

80 

60 

40 

20 

0 

0 

A2780 

500 

Energy exposure, J/cm 2 

1000 

A2780 

A2780/AD 

FIGURE 21.2 Immediate cell viability after ultrasound exposure decreases with increasing acoustic energy; 

cells were irradiated in suspensions by focused 1.1-MHz ultrasound. The MDR cells manifest higher susceptibility 

to the ultrasound action than their wild-type counterparts. (Adapted from Kamaev, P., and Rapoport, N., 

Therapeutic Ultrasound: 5th International Symposium on Therapeutic Ultrasound, Clement, G. T., McDannold, 

N. J., and Hynynen, K., Eds., American Institute of Physics, Melville, NY, pp. 543–547, 2006. With permission.) 

A2780/AD 

no DOX 

DOX 

FIGURE 21.3 Effect of DOX on the immediate viability of drug-sensitive and MDR ovarian carcinoma cells 

sonicated in suspensions by focused ultrasound; DOX concentration in the media is 10 mg/mL for the MDR 

cells and 20 mg/mL for the sensitive cells. Ultrasound conditions: frequency 1.1 MHz, pressure 0.55 MPa, 

duration 15 s, duty cycle 33%, room temperature. (Adapted from Kamaev, P., and Rapoport, N., Therapeutic 

Ultrasound: 5th International Symposium on Therapeutic Ultrasound, Clement, G. T., McDannold, N. J., and 

Hynynen, K., Eds., American Institute of Physics, Melville, NY, pp. 543–547, 2006. With permission.) 


1500

428 

drug-loaded nanoparticles. It was shown that ultrasound enhanced the extravasation of the macromolecular 

MRI contrast agent. 66 

21.3.3 ULTRASOUND-INDUCED DRUG RELEASE FROM POLYMERIC MICELLES 

Ultrasound was shown to release DOX from the unstabilized and stabilized Pluronic w micelles. 46,55 

A degree of drug release depended on a number of ultrasound parameters—frequency, power 

density, pulse length, and inter-pulse intervals. 46 Examples of drug release profiles under CW or 

pulsed ultrasound are presented in Figure 21.4. Measurements were based on a decrease of DOX 

fluorescence when drug was transferred from the hydrophobic environment of micelle cores into the 

aqueous environment. Measurements were performed in situ in the absence of cells. 

As indicated by fluorescence profiles, drug release under the action of ultrasound pulses was 

reversible and drug re-encapsulation proceeded during the inter-pulse intervals. Pulsed ultrasound 

induced alternating cycles of micelle perturbation and restoration. Kinetic parameters of the drug 

release and re-encapsulation indicated that both processes were controlled by motion of micelleforming 

macromolecules rather than diffusion of relatively small drug molecules. 55 

Low-frequency ultrasound was more effective in drug release from micelles than highfrequency 

ultrasound (Figure 21.5). Low-frequency ultrasound can penetrate deeper into the 

tissue than high-frequency ultrasound, but it does not allow sharp focusing. 50 Therefore, in vivo, 

high-frequency ultrasound could be used for irradiating small and relatively superficial tumors, 

while low-frequency ultrasound could be used for larger and deeper located tumors. 

21.3.4 ULTRASOUND-ENHANCED INTRACELLULAR DRUG UPTAKE 

Without ultrasound, drug encapsulation in polymer micelles results in a significantly reduced intracellular 

uptake. The degree of drug shielding depends on drug/micelle interaction, and is higher for 

Fluorescence intensity, Arb, Uniits 

(a) 

(b) 

(c) 

(d) 

Continuous wave 

Pulsed 

0.5s "on" : 0.5s "off" 

1s "on" : 1s "off" 

1s "on" : 3s "off" 

0 20 40 60 80 100 120 140 160 

Time (s) 


FIGURE 21.4 Doxorubicin (DOX) release profiles from 10% Pluronic micelles under continuous wave (CW) 

and pulsed 20 kHz ultrasound at a power density of 0.058 W/cm 2 at various durations of ultrasound pulses and 

inter-pulse intervals indicated in the figure. Measurements are based on the decrease of fluorescence intensity 

when DOX is transferred from the hydrophobic environment of micelles cores into the aqueous environment. 

(Adapted from Husseni, G. A. et al., J. Control. Release, 69, 43–52, 2000. With permission.) 


180


Degree of drug release, % 

14 

12 

10 

8 

6 

4 

2 

67 kHz 

1 MHz 

0 

0 2 4 

Power density, W/cm 

6 8 

2 

FIGURE 21.5 Effect of ultrasound frequency and power density on DOX release from 10% Pluronic P-105 

micelles. (Adapted from Marin, A. et al., J. Control. Release, 84, 39–47, 2002.) 

more lipophilic drugs. As an example, DOX uptake by ovarian carcinoma A2780 cells dropped by 

50% after encapsulation in 10% Pluronic w micelles. For a more hydrophobic analog of DOX called 

Ruboxyl, micellar encapsulation resulted in a decrease of the intracellular uptake by 66%. 

Drug shielding by micelles reduces systemic concentrations of toxic drugs, diminishing side 

effects. As an example, DOX encapsulation in PEO-b-poly(a-benzyl L-aspartate) (PEO–PBLA) 

micelles allowed administering high concentrations of DOX to tumor-bearing mice without 

toxicsideeffects. 67 The same results were manifested by micelle-encapsulated PTX 39–43 

(see Section 21.2.2). 

Ultrasound substantially enhanced the intracellular drug uptake from/with polymeric 

micelles. 45,50,56 Typical examples are shown in Figure 21.6. 

Based on the results presented above, a new modality of tumor-targeted chemotherapy has been 

suggested. According to this technology, micelle-encapsulated drugs are injected intravenously to 

tumor-bearing subjects. Drug-loaded micelles are long-circulating (Figure 21.7), which provides 

for their gradual accumulation in the tumor interstitium via the EPR effect (see below). At the time 

corresponding to a maximal micelle accumulation, ultrasound is applied locally to the tumor. Under 

the ultrasound action, drug is released from micelles in the tumor volume; the intracellular uptake 

of drug by tumor cells is enhanced due to ultrasound-induced plasma membrane perturbation. 

Important advantages of this technique are its noninvasive character, on-demand drug release, 

and a prospect for sensitization of drug-resistant tumors. 

Counts 

10 0 

600 

480 

360 

240 

120 

0 

1 

2 

10 1 

3 

4 

10 

Fluorescence, arb.u. 

2 

FIGURE 21.6 Fluorescence histogram of A2780 cells. (1) Control; (2)–(4): cells incubated for 30 min. with 

20 mg/mL DOX encapsulated in PEG–PLLA micelles. (2) No ultrasound; (3) sonication for 30 s by CW 

3-MHz ultrasound at 2.0 W/cm 2 ; (4) sonication for 30 s by 1-MHz ultrasound at 3.4 W/cm 2 and 33% duty 

cycle. (From Gao, Z. and Rapoport, N., unpublished results.) 


10 3 

10 4

430 

DOX Concentration in plasma, μg/ml 

10 

1 

0.1 

0.01 

0 4 8 12 

Time, h 

16 20 24 

In the following section, the results of the in vivo evaluation of the new therapeutic modality 

are discussed. 

21.4 IN VIVO EVALUATION OF THE DRUG/MICELLE/ULTRASOUND 

TUMOR-TARGETING MODALITY 

21.4.1 ULTRASOUND-INDUCED DRUG RELEASE FROM MICELLES IN VIVO 


FIGURE 21.7 Pharmacokinetics of DOX dissolved in PBS (squares) or encapsulated in stabilized Pluronic 

P-105 micelles (diamonds); mean values plus/minus standard deviations are presented, nZ4. (From Gao, Z. 

and Rapoport, N., unpublished results.) 

Drug release from micelles in vivo was confirmed in experiments described in studies by Gao, 

et al. 44 Ovarian carcinoma A2780 cells (2!10 6 ) were inoculated in peritoneal cavities of athymic 

(nu/nu) mice. The next day, mice were treated intraperitoneally by 3 mg/kg DOX that was either 

dissolved in PBS (designated “free DOX” hereafter) or encapsulated in 10% Pluronic w P-105 

micelles. Treatment was combined or not combined with the ultrasonic irradiation of the abdominal 

region of a mouse. Some groups of mice were treated by ultrasound without DOX. Tumor yields 

and animal survival rates were monitored. The results are presented in Table 21.1. 44 

Without DOX, no effect of ultrasound on the tumor staging was observed; the tumors were 

developed in 100% mice. Only a moderate effect of free DOX on the tumor development was 

observed without ultrasound, tumor yield being reduced by 30%. However, when DOX and ultrasound 

were combined, tumor yields were reduced by about 60% for both free and micellarencapsulated 

DOX, which was accompanied by a statistically significant increase of the average 

survival rates of corresponding animal groups. No differences were observed between the effects of 

free or micellar-encapsulated DOX, suggesting that ultrasound did, indeed, trigger DOX release 

from micelles in vivo. 

The absence of the ultrasound effect in the absence of DOX indicated that tumor cell death was 

associated with the ultrasound-enhanced cytotoxic action of DOX rather than mechanical or 

thermal action of ultrasound itself. 44 

21.4.2 BIODISTRIBUTION OF POLYMERIC MICELLES AND MICELLAR-ENCAPSULATED DRUGS 

The effect of ultrasound on biodistribution of polymeric micelles was studied using fluorescently 

labeled Pluronic w P-105 micelles. 47 The uptake of micelles by various organ cells was measured by 

flow cytometry. A significant advantage of this technique is that it does not require organ 



TABLE 21.1 

Tumor Yields and Survival Rates of Mice Inoculated on Day 0 with i.p. Injections of A2780 

Ovarian Carcinoma Cells 

Group 

Number of Animals 

in Group 

Number of Internal 

Tumors Tumor Yield (%) Survival Rate (Days) 

PBS 17 17 100 40.6G8.2 

PBSCUS 5 5 100 39.6G6.5 

P-105 empty micelles 13 13 100 51.9G21.6 

P-105 empty 

micelles/US 

13 13 100 47.2G9.0 

DOX/PBS a 

10 7 70 61.5G10 b 

DOX/PBS/US a 

5 2 40 77.3G24 

DOX/micelle/US a 

11 4 36 91G27 b 

a 

The mice were treated on days 1, 4, 7, and 11; DOX at 3 mg/kg in various delivery systems was i.p. injected. One hour after 

the DOX injection, some groups of animals were sonicated in the abdominal region for 30 s by 1-MHz ultrasound at a power 

density of 1.2 W/cm 2 . By day 25, all control mice developed tumors. The last three groups marked by “a” were treated on 

days 1, 26, and 35. 

b pZ.01. 

Source: Gao, Z., Fain, H., and Rapoport, N., J. Control. Release, 102, 203–221, 2005. 

homogenization, thus providing for direct measurements of carrier or drug concentrations inside the 

cells, i.e., at the site of action. 

A 100-mL solution of fluorescently labeled stabilized Pluronic w micelles was injected intravenously 

through the tail vein to ovarian carcinoma bearing nu/nu mice. At various time points 

ranging from 2 to 12 h after the i.v. injection, mice were sonicated for 30 s by 1-MHz or 3-MHz 

ultrasound. Ten minutes after the ultrasound application, mice were sacrificed, tumor and other 

organs excised, dried with filter paper, weighed, and sliced in trypsin. The individual cells of 

various organs were fixed by 2.5% glutaraldehyde. Fluorescence of Pluronic w P-105 molecules 

internalized by the cells of various organs was measured by flow cytometry. 

Without ultrasound, a multimodal distribution of tumor cell fluorescence was observed for 

the intravenous injections of the fluorescently labeled, stabilized Pluronic w micelles. A significant 

fraction of the cells manifested a relatively low fluorescence (Figure 21.8, regular line). Fluorescence 

intensity of other organ cells was at the level of the autofluorescence of corresponding 

cells. These data indicated low efficiency of the intracellular uptake of intact Pluronic w micelles. 

A short (30-s ultrasound exposure) tumor sonication strongly enhanced the intracellular uptake of 

Pluronic w molecules by the tumor cells (Figure 21.8, bold line), suggesting that before sonication, 

micelles were effectively accumulated in the tumor interstitium but were not effectively internalized 

by the tumor cells. Tumor sonication not only substantially enhanced micelle internalization, 

but also resulted in a much more uniform fluorescence distribution over the tumor cell population, 

suggesting an enhanced micelle diffusion over the tumor volume. 47 

The biodistribution of micellar-encapsulated DOX closely followed that of a micellar carrier. 44 

For micellar-encapsulated DOX without ultrasound, DOX uptake by the tumor cells was relatively 

low. A local tumor sonication by 1-MHz ultrasound dramatically increased the intracellular uptake 

of drug by the tumor cells (Figure 21.9). A similar effect was observed for PEG–PBLA micelles 

(Figure 21.10). 44 Somewhat enhanced DOX uptake by the cells of other organs was observed for 

1-MHz ultrasound (data not shown). This was associated with a deep penetration of 1-MHz ultrasound 

in the mouse body, resulting in ultrasound scattering at various internal interfaces and/or 


432 

Counts 

10 0 

150 

120 

90 

60 

30 

0 

reflection from the skin/air interface at the site opposite to the transducer. This unwanted effect was 

eliminated when tumor was irradiated by a more localized 3-MHz ultrasound (Figure 21.10). 

21.4.3 EFFECT OF THE TIME OF ULTRASOUND APPLICATION 

Measuring micelle accumulation in the tumor cells for various times of ultrasound application after 

drug injection enabled the selection of optimal ultrasound application times. For the stabilized 

Pluronic w micelles, the optimal time of ultrasound application was found to be between four and 

eight hours after injection. 44,47 The results were interpreted as follows. At shorter times, a large 

number of micelles were still circulating in the vasculature while a lower concentration of micelles 

was accumulated in the tumor interstitial space. At longer times, some clearance of micelles from 

the tumor volume could already have occurred. The optimal time of ultrasound application is 

expected to depend on the pharmacokinetics of a particular delivery system. 

21.4.4 INHIBITION OF TUMOR GROWTH USING MICELLE/ULTRASOUND DRUG DELIVERY 

21.4.4.1 Drug-Sensitive Tumors 

1 

10 1 

10 


2 

FIGURE 21.8 Fluorescence histogram of the tumor cells in (1) unsonicated, and (2) sonicated mouse; 5% 

stabilized Pluronic micelles comprising 0.1% fluorescently labeled Pluronic P-105 were injected intravenously 

to two subcutaneous A2780 tumor bearing mice; 4 h after the injection, the tumor of one mouse was sonicated 

for a total of 60 s by 1-MHz ultrasound at 3.4 W/cm 2 power density at 50% duty cycle (two 30-s sonications 

were performed, with a 30-s interval between the sonications for cooling the ultrasound probe); both mice were 

sacrificed 10 min after the completion of sonication. (Modified from Gao, Z., Fain, H., and Rapoport, N., Mol. 

Pharmaceutics, 1, 317–330, 2004.) 

Targeting of micellar-encapsulated drugs to tumor cells and uniform drug distribution in the tumor 

volume result in a significant suppression of the growth of subcutaneous ovarian carcinoma tumors 

Counts 

10 0 

100 

80 

60 

40 

20 

0 

1 

10 1 

2 

10 


2 

FIGURE 21.9 Fluorescence histograms of the tumor cells in (1) nonsonicated, and (2) sonicated mice injected 

i.v. with DOX (6 mg/kg) encapsulated in mixed micelles. Mice were sonicated for 30 s by 1-MHz CW 

ultrasound at a power density of 1.7 W/cm 2 applied 8 h after the drug injection; nonsonicated and sonicated 

mice were sacrificed 10 min later. (Modified from Gao, Z., Fain, H., and Rapoport, N., J. Control. Release, 

102, 203–221, 2005.) 



10 3 

10 3 

2 

10 4 

10 4


Counts 

(a) 

(b) 

(c) 

(d) 

10 0 

inoculated in nu/nu mice (Figure 21.11). 44 However, the success of the treatment depended 

dramatically on the stage of tumor progression at the start of the treatment. When the initial 

tumor volume at the start of the treatment exceeded 250 mm 3 , none of the treatment modalities 

was effective in suppressing tumor growth, which emphasizes a need for early tumor detection. 

21.4.4.2 Drug Resistant Tumors 

10 1 

1 

We have attained substantial success in treating multidrug resistant, poorly differentiated MCF- 

7/ADmt breast cancer tumors by micellar-encapsulated PTX combined with ultrasonic tumor irradiation. 

The MCF7/AD mt tumor cells expressed the MDR1 gene 25 and were highly resistant to 

treatment by clinical PTX formulation, Taxol w (PTX injection). However, tumor growth was 

substantially suppressed when animals were treated by intravenous injections of a micellar PTX 

formulation in conjunction with local tumor sonication (Figure 21.12) (the micellar formulation of 

PTX, Genexol w was kindly donated by Samyang Pharmaceuticals Inc. (Korea)). 

In vitro without ultrasound, the intracellular uptake of PTX from the micellar formulation was 

significantly lower than that with PTX injection, which can be advantageous when used in vivo for 

preventing unwanted drug interactions with healthy cells. Under ultrasound, drug uptake by the 

MCF7/AD mt cells from micellar PTX formulation was increased more than 20-fold, which resulted 

in the inhibition of cellular proliferation by nearly 90%. 

10 2 

Fluorescence,arb.u. 

2 

10 3 

Kidney 

Spleen 

FIGURE 21.10 Fluorescence histograms of (a) tumor, (b) kidney, (c) spleen, and (d) heart cells of (1) 

nonsonicated (regular lines), and (2) sonicated (bold lines) mice injected with DOX (6 mg/kg) encapsulated 

in PEG–PBLA micelles. Mice were sonicated for 30 s by 3-MHz CW ultrasound 12 h after the drug injection at 

a power density of 1.8 W/cm 2 . (Adapted from Gao, Z., Fain, H., and Rapoport, N., J. Control. Release, 102, 

203–221, 2005.) 


Heart 

10 4

434 

Tumor volume, mm 3 

2500 

2000 

1500 

1000 

500 

0 

0 10 20 30 40 

Days after first treatment 

control 

DOX/PBS 

DOX/micelles 

DOX/micelles/US 

FIGURE 21.11 Growth curves of s.c. A2780 ovarian carcinoma tumors inoculated in female nu/nu mice; 

DOX (3 mg/kg) was either dissolved in PBS (positive control) or encapsulated in mixed micelles; drug 

injections were combined or not combined with tumor sonication. Treatments were initiated when the 

tumor volume reached 75–125 mm 3 ; three consecutive treatments were applied on days 1, 3, and 5. Ultrasound 

parameters: frequency, 1 MHz; power density, 3.4 W/cm 2 ; duty cycle 50%; duration, 30 s; time of application: 

4 h after the i.v. injection of the drug. (Adapted from Gao, Z., Fain, H., and Rapoport, N., J. Control. Release, 

102, 203–221, 2005.) 

The in vivo results are shown in Figure 21.12. As could be expected, without ultrasound, 

clinical PTX and Genexol w both manifested low efficacy; in contrast, injections of Genexol w 

combined with ultrasonic tumor irradiation resulted in a complete tumor resolution. Most importantly, 

the tumors treated by Genexol w /ultrasound did not metastasize, whereas tumors treated 

without ultrasound manifested a high rate of metastasizing into lymph nodes and/or spleen. 

Total Tumor Volume (mm 3 ) 

550 

500 

Control 

Clinical paclitaxel 

Genexol 

450 Genexol + US 

400 

350 

300 

250 

200 

150 

100 

50 

0 

0 20 40 

Time (days) 

60 80 


FIGURE 21.12 Growth curves of subcutaneous MCF7/ADmt tumors inoculated in female nu/nu mice; bolus 

injections of a clinical paclitaxel formulation Paclitaxel injections or a micellar formulation Genexol w 

provided by Samyang Pharmaceuticals, Inc. (Korea) were administered twice a week for two weeks intravenously 

through the tail vein at a dose rate of 13 mg/kg/day. (From Howard, B., Gao, Z., and Rapoport, N., Am. 

J. Drug Deliv., in press.) 



The aqueous base of the drug formulation, reduced systemic toxicity, potentials for tumor 

targeting, and on-demand drug release triggered by ultrasound are important additional advantages 

of micellar formulations of PTX. 

21.4.4.3 Poorly Vascularized Tumors 

A different type of drug resistance was observed for subcutaneously grown colon cancer HCT116 

tumors. These tumors demonstrated very loose capillary networks and large necrotic areas filled 

with ascetic fluid. HCT116 colon cancer tumors manifested strong resistance to the intravenous 

drug injections independent on the delivery modality (Figure 21.13, upper set of curves). The 

resistance was caused by poor vascularization of the tumor volume, resulting in insufficient drug 

delivery to tumor cells. 

Poor tumor vascularization represents a unique type of tumor resistance to chemotherapy. Upon 

intravenous drug injections, avascular tumor regions receive insufficient levels of chemotherapeutic 

agents; cancerous cells in these regions remain viable and promote tumor growth. This type of 

resistance has not been widely discussed in the literature. 

In our laboratory, successful treatment of HCT116 tumors was achieved by direct intratumoral 

injections of micellar-encapsulated DOX (Figure 21.13, lower set of curves). The encapsulated 

drug was effectively retained in the tumor volume, with very low, if any, accumulation in the cells 

of other organs. In contrast, free DOX quickly leaked from the tumor tissue, and a short time after 

the intratumoral injection of free drug, DOX was already found in the heart cells and those of other 

organs, causing associated toxicity. 

With the intratumoral injections of micellar encapsulated DOX, effective tumor suppression 

was observed even without ultrasound (Figure 21.13). We hypothesize that a high hydrostatic 

pressure generated by intratumoral injections forced drug-loaded micelles through cell membranes. 

For therapy of poorly vascularized tumors, intratumoral injections of micellar-encapsulated 

drugs present a viable alternative to inefficient intravenous injections. 

The advantages of intratumoral injections were also observed for highly vascularized A2780 

tumors. The beneficial effect of ultrasound was revealed in drug biodistribution experiments. For 

the intravenous injections of micellar-encapsulated DOX, tumor sonication resulted in a 

Tumor volume, mm 

2500 

2000 

1500 

1000 

500 

0 

0 20 

40 

Days after first treatment 

Intravenous 

Injections 

Intratumaoral 

Injections 

60 80 100 

FIGURE 21.13 Growth curves of HCT116 tumors upon the intravenous (upper group of curves) or intratumoral 

(lower group of curves) injections of DOX at a dose rate of 3 mg/kg. Intravenous injections: filled 

diamonds—control, open triangles—DOX/PBS, open squares—DOX/mixed micelles/ultrasound; Intratumoral 

injections: open circles—DOX/mixed micelles, closed circles—DOX/mixed micelles/ultrasound. 

(From Gao, Z. and Rapoport, N., unpublished results.) 


436 

dramatically increased intracellular drug uptake; for the intratumoral injections of the same formulation, 

the main effect of ultrasound was related to a much more uniform drug distribution over the 

tumor cell population (Figure 21.14). 

Intratumoral chemotherapy by nanoparticle encapsulated drugs could be advantageous for 

tumors with well-defined primary lesions such as breast, colorectal, prostate, and skin cancers. 

These procedures may be performed noninvasively or minimally invasively through intraluminal, 

laparoscopic or percutaneous means under the guidance of an appropriate imaging modality such as 

MRI, CT, or ultrasound. During the last decade, the interest in intratumoral chemotherapy has 

significantly increased. 68 

21.4.5 MECHANISMS OF MICELLE/ULTRASOUND BIOEFFECTS 

Ultrasound irradiation may play multiple roles in tumor therapy. Ultrasound may (1) induce thermal 

effects, (2) enhance extravasation of drug-loaded micelles into the tumor interstitium, (3) enhance 

drug diffusion through the tumor interstitium, (4) trigger drug release from micelles in the tumor 

volume, and (5) enhance the intracellular uptake of both released and encapsulated drug at the 

irradiation site. Though the relative importance of these factors remains to be further explored, 

some conclusions can be drawn based on the results presented above. 

Drug release from perturbed micelles under the action of ultrasound is undoubtedly an important 

factor in the ultrasound-enhanced drug uptake by the tumor cells from micellar carriers. 

Ultrasound can also enhance the intracellular uptake of intact drug-loaded micelles due to cell 

membrane perturbation. Another important component of the ultrasound effect is associated with 

the ultrasound-enhanced diffusion of free or micellar-encapsulated drug over tumor tissue. 

In summary, drug encapsulation in polymeric micelles combined with local ultrasonic tumor 

irradiation results in effective tumor targeting and suppression of growth of drug-sensitive and 

multidrug resistant tumors. The effect is associated with the ultrasound-enhanced cytotoxic action 

of passively targeted or intratumorally injected micellar-encapsulated chemotherapeutic agents. 

(a) 

(b) 

Counts 

Counts 

10 0 

10 0 

1 

10 1 

10 1 

10 2 

10 2 


10 3 

1 2 

FIGURE 21.14 Fluorescence histograms of A2780 tumor cells in (a) unsonicated, and (b) sonicated mouse; 

DOX encapsulated in stabilized Pluronic micelles was injected at a dose of at 6 mg/kg (1) intravenously, and 

(2) intratumorally; sonicated tumors were irradiated for 30 s by 1-MHz ultrasound at 3.4 W/cm 2 power density 

and 50% duty cycle 4 h after the drug injection. (From Gao, Z. and Rapoport, N., unpublished results.) 


2 


10 3 

10 4 

10 4


This allows on-demand drug delivery and control of systemic, local, and subcellular 

drug distribution. 

21.5 COMBINING ULTRASOUND IMAGING AND TREATMENT 

Ultrasonic tumor treatment requires tumor imaging prior to therapy. Using ultrasound for both 

tumor imaging and treatment is especially attractive. Development of ultrasound contrast agents 

and concomitant use of harmonic ultrasound imaging has recently made possible imaging of liver, 69 

breast, 70 and prostate tumors. 71 Ultrasound-stimulated acoustic emission of microbubbles was used 

for color Doppler imaging of liver metastases. 72 

Some dual ultrasonic imaging/treatment techniques are already in clinical practice, others are in 

the investigational stage. Examples include Sonablate w for prostate cancer treatment (Focus 

Surgery, Inc., Indianapolis, Indiana) and a miniature ultrasound-based dual-modality device for 

image guided soft tissue ablation (Guided Therapy Systems, Mesa, Arizona). 

With dual-modality ultrasound instruments already on the market, developing dual-modality 

ultrasound imaging/drug carrier systemswould be very timely. 

21.5.1 DUAL-MODALITY CONTRAST AGENTS/DRUG CARRIER SYSTEMS 

Ultrasonically enhanced site-specific drug delivery using acoustically activated drug-loaded microbubbles 

was suggested by Ungar 73 and Ferrara; 74–77 these authors developed echogenic 

microbubbles stabilized by lipid bilayers. Drug was incorporated in the oil layer located on the 

microbubble surface. An effective ultrasound-triggered transfer of a solute from the modified 

lipospheres to the tumor cells has been demonstrated. Polymeric ultrasound contrast agents with 

targeting potentials have also been explored by the Wheatley group. 78–81 

Nano/microbubbles oscillate in the ultrasound field and can be cavitated for site-specific localized 

delivery of chemotherapeutic agents; they concentrate ultrasound energy and may be used for 

enhancing drug delivery from micelles and other nanoparticles. Nano/microemulsions and microbubbles 

are efficient reflectors of ultrasound energy, which may be used for contrast-enhanced 

ultrasound imaging of tumor neovasculature. 

We have recently developed a new class of these agents that combine diagnostic and therapeutic 

properties and can be used for ultrasonically enhanced drug delivery and tumor imaging. Our 

microbubbles differ in a number of ways from the products on the market or those under development 

by other groups: 

† The microbubbles are produced in situ upon injection of specially designed microemulsions. 

† The microbubbles have strong walls produced by biodegradable diblock copolymers; the 

walls stabilize microbubbles for ultrasound imaging. 

† The diblock copolymer not only forms microbubble walls but also forms polymeric 

micelles that effectively encapsulate chemotherapeutic agents and act as drug carriers. 

† Upon injection, a localized ultrasonic irradiation of the tumor in the presence of microbubbles 

provides for effective intracellular drug uptake by the tumor cells, which results 

in inhibition of tumor growth. 

Echogenic nano/microbubbles designed to interact specifically with neovascular epithelium 

may allow ultrasound imaging of small metastases. To ensure molecular targeting, the nano/microparticle 

surface should be decorated by site-specific ligands. In the future, we anticipate conjugating 

block copolymer molecules with vectors that would target micelles and nano/microbubbles to the 

receptors that are overexpressed on the endothelial lining of tumor neovasculature. 


438 


In animal models, drug encapsulation within polymeric micelles combined with ultrasonic tumor 

irradiation results in effective tumor targeting and inhibition of tumor growth. The effect is based on 

ultrasonically triggered localized drug release from micelles and enhanced intracellular drug 

uptake. This technique offers prospects for treating multidrug resistant tumors that fail conventional 

chemotherapy. Dual-modality drug delivery/ultrasound contrast agent systems used with dualmodality 

ultrasound imaging/treatment instruments are expected to allow precisely controlled, 

ultrasound-enhanced tumor chemotherapy in a clinical environment. 

REFERENCES 


1. Kwon, G. S. and Kataoka, K., Block copolymer micelles as long circulating drug vehicles, Drug 

Deliv. Rev., 16, 295, 1995. 

2. Aliabadi, H. and Lavasanifar, A., Polymeric micelles for drug delivery, Expert Opin. Drug Deliv., 3, 

139, 2006. 




4. Zhang, X., Burt, H. M., and Von Hoff, D., An investigation of the antitumor activity and biodistribution 

of polymeric micellar paclitaxel, Cancer Chemother. Pharmacol., 40, 81, 1997. 


micellear carriers of taxol, Int. J. Pharm., 132, 195, 1996. 

6. Zhang, X., Burt, H. M., and Mangold, G., Anti-tumor efficacy and biodistribution of intravenous 

polymeric micellar paclitaxel, Anticancer Drug, 8, 696, 1997. 

7. Adams, M. L., Andes, D. R., and Kwon, G. S., Amphotericin B encapsulated in micelles based on 

poly(ethylene oxide)-b-poly(L-amino acid) derivatives exerts reduced in vitro hemolysis but maintains 

potent in vivo antifungal activity, Biomacromolecules, 4, 750, 2003. 


encapsulated by poly(ethylene oxide)-b-poly(N-hexyl-L-aspartamide)-acyl conjugate micelles: 

Effects of acyl chain length, J. Control. Release, 87, 23, 2003. 

9. Leung, S., Jackson, J., Miyake, H., Burth, H., and Gleave, M. E., Polymeric micellar paclitaxel 

phosphorylates Bcl-2 and induces apopototic regression of androgen-independent LNCaP prostate 

tumors, Prostate, 44, 156, 2000. 

10. Shin, I. G. et al., Methoxy poly(ethylene glycol)/epsilon-caprolactone amphiphilic block copolymeric 

micelle containing indomethacin. I. Preparation and characterization, J. Control. Release, 51, 1, 1998. 

11. Kim, S. C. et al., In vivo evaluation of polymeric micellar paclitaxel formulation: Toxicity and 

efficacy, J. Control. Release, 72, 191, 2001. 

12. Benahmed, A., Ranger, M., and Leroux, J. C., Novel polymeric micelles based on the amphiphilic 

diblock copolymer poly(N-vinyl-2-pyrrolidone)-b-poly(D,L-lactide), Pharm. Res., 18, 323, 2001. 

13. Yu, B. G. et al., Polymeric micelles for drug delivery: Solubilization and haemolytic activity of 

amphotericin B, J. Control. Release, 53, 131, 1998. 

14. Lavasanifar, A., Samuel, J., and Kwon, G. S., Micelles of poly(ethylene oxide)-b-poly(N-alkyl stearate 

L-aspartamide): Synthetic analogues of lipoproteins for drug delivery, J. Biomed. Mater. Res., 52, 

831, 2000. 

15. Piskin, E., Kaitian, X., Denkbas, E. B., and Kucukyavuz, Z., Novel PDLLA/PEG copolymer micelles 

as drug carriers, J. Biomater. Sci. Polym. Ed., 7, 359, 1995. 

16. Francis, M. F., Piredda, M., and Winnik, F. M., Solubilization of poorly water soluble drugs in 

micelles of hydrophobically modified hydroxypropylcellulose copolymers, J. Control. Release, 93, 

59, 2003. 

17. Jette, K. K. et al., Preparation and drug loading of poly(ethylene glycol)-b-poly(epsilon-caprolactone) 

micelles through the evaporation of a cosolvent azeotrope, Pharm. Res., 21, 1184, 2004. 

18. Nakanishi, T. et al., Development of the polymer micelle carrier system for doxorubicin, J. Control. 

Release, 74, 295, 2001. 



19. Rapoport, N., Stabilization and acoustic activation of pluronic micelles for tumor-targeted drug 

delivery, Colloids Surf. B Biointerfaces, 3, 93, 1999. 

20. Shuai, X. et al., Core-cross-linked polymeric micelles as paclitaxel carriers, Bioconj. Chem., 15, 441, 

2004. 

21. Lee, J., Cho, E. C., and Cho, K., Incorporation and release behavior of hydrophobic drug in functionalized 

poly(D,L-lactide)-b-poly(ethylene oxide) micelles, J. Control. Release, 94, 323, 2004. 

22. Matsumura, Y. et al., Phase I clinical trial and pharmacokinetic evaluation of NK911, a micelleencapsulated 

doxorubicin, Br. J. Cancer, 91, 1775, 2004. 

23. Alakhov, V. et al., Hypersensitization of multidrug resistant human ovarian carcinoma cells by 

pluronic P85 block copolymer, Bioconjug. Chem., 7, 209, 1996. 

24. Batrakova, E. V. et al., Anthracycline antibiotics non-covalently incorporated into the block copolymer 

micelles: In vivo evaluation of anti-cancer activity, Br. J. Cancer, 74, 1545, 1996. 

25. Venne, A. et al., Hypersensitizing effect of pluronic L61 on cytotoxic activity, transport, and subcellular 

distribution of doxorubicin in multiple drug-resistant cells, Cancer Res., 56, 3626, 1996. 

26. Batrakova, E. V. et al., Effects of pluronic block copolymers on drug absorption in Caco-2 cell 

monolayers, Pharm. Res., 850, 1998. 

27. Batrakova, E. V. et al., Effects of pluronic P85 unimers and micelles on drug permeability in polarized 

BBMEC and Caco-2 cells, Pharm. Res., 15, 1525, 1998. 

28. Miller, D. W., Batrakova, E. V., and Kabanov, A. V., Inhibition of multidrug resistance-associated 

protein (MRP) functional activity with pluronic block copolymers, Pharm. Res., 16, 396, 1999. 

29. Batrakova, E. et al., Fundamental relationships between the composition of pluronic block copolymers 

and their hypersensitization effect in MDR cancer cells, Pharm. Res., 16, 1373, 1999. 

30. Alakhov, V. et al., Block copolymer-based formulation of doxorubicin. From cell screen to clinical 

trials, Colloids Surf. B Biointerfaces, 16, 113, 1999. 

31. Batrakova, E. V. et al., Optimal structure requirements for pluronic block copolymers in modifying 

P-glycoprotein drug efflux transporter activity in bovine brain microvessel endothelial cells, 

J. Pharmacol. Exp. Ther., 304, 845, 2003. 

32. Kabanov, A. V., Batrakova, E. V., and Alakhov, V. Y., An essential relationship between ATP 

depletion and chemosensitizing activity of Pluronic block copolymers, J. Control. Release, 91, 75, 

2003. 

33. Venne, A. et al., Hypersensitizing effect of pluronic L61 on cytotoxic activity, transport, and subcellular 

distribution of doxorubicin in multiple drug-resistant cells, Cancer Res., 56, 3626, 1996. 

34. Rapoport, N. et al., Intracellular uptake and trafficking of Pluronic micelles in drug-sensitive and 

MDR cells: Effect on the intracellular drug localization, J. Pharm. Sci., 91, 157, 2002. 

35. Danson, S. et al., Phase I dose escalation and pharmacokinetic study of pluronic polymer-bound 

doxorubicin (SP1049C) in patients with advanced cancer, Br. J. Cancer, 90, 2085, 2004. 

36. Danson, S. et al., Phase I dose escalation and pharmacokinetic study of pluronic polymer-bound 

doxorubicin (SP1049C) in patients with advanced cancer, Br. J. Cancer, 90, 2085, 2004. 

37. Valle, J. W. et al., A phase II, window study of SP1049C as first-line therapy in inoperable metastatic 

adenocarcinoma of the esophagus, J. Clin. Oncol., 22, 4195, 2004. 

38. Gelderblom, H. et al., Cremophor EL: The drawbacks and advantages of vehicle selection for drug 

formulation, Eur. J. Cancer, 37, 1590, 2001. 

39. Kim, S. C. et al., In vivo evaluation of polymeric micellar paclitaxel formulation: Toxicity and 

efficacy, J. Control. Release, 72, 191, 2001. 

40. Kim, T. Y. et al., Phase I and pharmacokinetic study of Genexol-PM, a cremophor-free, polymeric 

micelle-formulated paclitaxel, in patients with advanced malignancies, Clin. Cancer Res., 10, 3708, 

2004. 

41. Fournier, E. et al., A novel one-step drug-loading procedure for water-soluble amphiphilic nanocarriers, 

Pharm. Res., 21, 962, 2004. 

42. Le Garrec, D. et al., Poly(N-vinylpyrrolidone)-b-poly(D,L-lactide) as a new polymeric solubilizer for 

hydrophobic anticancer drugs: In vitro and in vivo evaluation, J. Control. Release, 99, 83, 2004. 

43. Hamaguchi, T. et al., NK105, a paclitaxel-incorporating micellar nanoparticle formulation, can extend 

in vivo antitumour activity and reduce the neurotoxicity of paclitaxel, Br. J. Cancer, 92, 1240, 2005. 


drug and ultrasound, J. Control. Release, 102, 203, 2005. 


440 


45. Marin, A., Muniruzzaman, M., and Rapoport, N., Mechanism of the ultrasound activation of micellar 

drug delivery, J. Control. Release, 75, 69, 2001. 

46. Husseini, G. A. et al., Factors affecting acoustically triggered release of drugs from polymeric 

micelles, J. Control. Release, 69, 43, 2000. 

47. Gao, Z., Fain, H. D., and Rapoport, N., Ultrasound-enhanced tumor targeting of polymeric micellar 

drug carriers, Mol. Pharm., 1, 317, 2004. 

48. Rapoport, N., Tumor targeting by polymeric assemblies and ultrasound activation, In Kentus Books 

MML Series, Arshadi, R. and Kono, K., Eds., Vol. 8, Kentus Books, London, pp. 305–362, 2005. 

49. Rapoport, N., Combined cancer therapy by micellar-encapsulated drug and ultrasound, Int. J. Pharm., 

277, 155, 2004. 

50. Rapoport, N., Marin, A., and Christensen, D. A., Ultrasound-activated drug delivery, Drug Deliv. Syst. 

Sci., 2, 37, 2002. 

51. Rapoport, N., Factors affecting ultrasound interactions with polymeric micelles and viable cells, In 

Carrier-Based Drug Delivery, ACS Symposium Series, Swenson, S., Ed., Vol. 879, Chapter 12, 

American Chemical Society, Washington, DC, 2004. 

52. Rapoport, N., Marin, A., Muniruzzaman, M., and Christensen, D., Controlled drug delivery to drugsencitive 

and multidrug resistant cells: Effects of pluronic micelles and ultrasound, In Advances in 

Controlled Drug Delivery, ACS Symposium Book Series, Dinh, S. M. and Liu, P., Eds., American 

Chemical Society, Washington, DC, pp. 85–101, 2003. 

53. Rapoport, N. et al., Drug delivery in polymeric micelles: From in vitro to in vivo, J. Control. Release, 

91, 85, 2003. 

54. Rapoport, N. Y. et al., Ultrasound-triggered drug targeting of tumors in vitro and in vivo, Ultrasonics, 

42, 943, 2004. 

55. Husseini, G. A. et al., Kinetics of ultrasonic release of Doxorubicin from Pluronic P-105 micelles, 

Colloids Surf. B Biointerfaces, 24, 253, 2002. 

56. Marin, A. et al., Drug delivery in pluronic micelles: Effect of high-frequency ultrasound on drug 

release from micelles and intracellular uptake, J. Control. Release, 84, 39, 2002. 

57. Deng, C. X. et al., Ultrasound-induced cell membrane porosity, Ultrasound Med. Biol., 30, 519, 2004. 

58. Honda, H. et al., Role of intracellular calcium ions and reactive oxygen species in apoptosis induced 

by ultrasound, Ultrasound Med. Biol., 30, 683, 2004. 

59. Pan, H. et al., Study of sonoporation dynamics affected by ultrasound duty cycle, Ultrasound Med. 

Biol., 31, 849, 2005. 

60. Miller, D. L. and Quddus, J., Sonoporation of monolayer cells by diagnostic ultrasound activation of 

contrast-agent gas bodies, Ultrasound Med. Biol., 26, 661, 2000. 

61. Kamaev, P. P. and Rapoport, N. Y., Effect of anticancer drug on the cell sensitivity to ultrasound in 

vitro and in vivo, AIP Conf. Proc., 829, 543, 2006. 

62. Li, P., Armstrong, W. F., and Miller, D. L., Impact of myocardial contrast echocardiography on 

vascular permeability: Comparison of three different contrast agents, Ultrasound Med. Biol., 30, 

83, 2004. 

63. Li, P. et al., Impact of myocardial contrast echocardiography on vascular permeability: An in vivo 

dose response study of delivery mode, pressure amplitude and contrast dose, Ultrasound Med. Biol., 

29, 1341, 2003. 

64. Miller, D. L. et al., Histological characterization of microlesions induced by myocardial contrast 

echocardiography, Echocardiography, 22, 25, 2005. 

65. Wible, J. H. et al., Microbubbles induce renal hemorrhage when exposed to diagnostic ultrasound in 

anesthetized rats, Ultrasound Med. Biol., 28, 1535, 2002. 

66. Bednarski, M. D. et al., In vivo target-specific delivery of macromolecular agents with MR-guided 

focused ultrasound, Radiology, 204, 263, 1997. 

67. Kataoka, K. et al., Doxorubicin-loaded poly(ethylene glycol)-poly(beta-benzyl-L-aspartate) 

copolymer micelles: Their pharmaceutical characteristics and biological significance, J. Control. 

Release, 64, 143, 2000. 

68. Goldberg, E. P. et al., Intratumoral cancer chemotherapy and immunotherapy: Opportunities for 

non-systemic preoperative drug delivery, J. Pharm. Pharmacol., 54, 159, 2002. 



69. Wilson, S. R. et al., Harmonic hepatic US with microbubble contrast agent: Initial experience showing 

improved characterization of hemangioma, hepatocellular carcinoma, and metastasis, Radiology, 215, 

153, 2000. 

70. Kedar, R. P. et al., Microbubble contrast agent for color Doppler US: Effect on breast masses. Work in 

progress, Radiology, 198, 679, 1996. 

71. Halpern, E. J., Rosenberg, M., and Gomella, L. G., Prostate cancer: Contrast-enhanced ultrasound for 

detection, Radiology, 219, 219, 2001. 

72. Bauer, A. et al., Ultrasonic imaging of organ perfusion with SH U 563A, Acad. Radiol., 9(Suppl 1), 

S46, 2002. 

73. Unger, E. C. et al., Therapeutic applications of lipid-coated microbubbles, Adv. Drug Deliv. Rev., 56, 

1291, 2004. 

74. Shortencarier, M. J. et al., A method for radiation-force localized drug delivery using gas-filled 

lipospheres, IEEE Trans. Ultrason. Ferroelectr. Freq. Control, 51, 822, 2004. 

75. Bloch, S. H., Dayton, P. A., and Ferrara, K. W., Targeted imaging using ultrasound contrast agents. 

Progess and opportunities for clinical and research applications, IEEE Eng. Med. Biol. Mag., 23, 18, 

2004. 

76. Bloch, S. H. et al., The effect of size on the acoustic response of polymer-shelled contrast agents, 

Ultrasound Med. Biol., 31, 439, 2005. 

77. Dayton, P. A. and Ferrara, K. W., Targeted imaging using ultrasound, J. Magn. Reson. Imaging, 16, 

362, 2002. 

78. Lathia, J. D., Leodore, L., and Wheatley, M. A., Polymeric contrast agent with targeting potential, 

Ultrasonics, 42, 763, 2004. 

79. El-Sherif, D. M. et al., Ultrasound degradation of novel polymer contrast agents, J. Biomed. Mater. 

Res. A, 68, 71, 2004. 

80. El-Sherif, D. M. and Wheatley, M. A., Development of a novel method for synthesis of a polymeric 

ultrasound contrast agent, J. Biomed. Mater. Res. A, 66, 347, 2003. 

81. Basude, R. and Wheatley, M. A., Generation of ultraharmonics in surfactant based ultrasound contrast 

agents: Use and advantages, Ultrasonics, 39, 437, 2001. 


22 

CONTENTS 

Polymeric Micelles Targeting 

Tumor pH 

Eun Seong Lee and You Han Bae 

22.1 Introduction ....................................................................................................................... 443 

22.2 Acidic Tumor pH e ............................................................................................................. 444 

22.3 Multidrug Resistance and Cancer ..................................................................................... 444 

22.4 PolyHis-PEG-Based pH-Sensitive Micelles: Preparation and Characteristics ................ 445 

22.4.1 Synthesis .............................................................................................................. 445 

22.4.2 PolyHis/PEG Copolymer Micelles ..................................................................... 446 

22.4.3 Mixed Micelles of PolyHis/PEG and PLLA/PEG.............................................. 448 

22.5 Tumor pHe Targeting ........................................................................................................ 451 

22.6 Early Endosomal pH Targeting for Reversal of MDR .................................................... 455 

22.7 Conclusions ....................................................................................................................... 460 

References..................................................................................................................................... 461 


The dose–response curves for cell cycle-phase nonspecific anti-cancer drugs, such as doxorubicin, 

have been constructed from in vitro experimental results. The survival fraction of both wild-type 

sensitive and resistant cancer cells decreases in a sigmoidal pattern as a function of log of dose size 

(the survival fraction of the cells remains the same as the untreated control at low concentrations, 

starts to sharply decrease above a specific concentration and reaches a plateau at high concentrations). 

The inflection point of the curves is drug specific and is shifted to a lower concentration 

with an increasing duration of drug exposure. The inflection point for the resistant cells occurs at a 

higher concentration region and shows a more gradual decrease in viability with increasing drug 

concentration. 1,2 Overall, the experimental results and theoretical fitting of these results give us a 

simple implication that high dose exposure for a longer duration results in effective destruction. 

Two classical nanosized anti-cancer drug carriers are liposomes and polymeric micelles. Each 

carrier system shows advantages as well as disadvantages over the other. For instance, liposomal 

carriers show continuous leakage of an entrapped drug with instability during storage and circulation. 

Polymeric micelles demonstrate much improved stability, but they may still dissociate upon 

dilution after injection and interact with blood components. 3,4 

Considering the dose–response relationship, it was suggested that future improvement in carrier 

design, providing a triggerable mechanism for drug release upon reaching the tumor sites, could be 

the most efficacious delivery strategy. 5 Various approaches for externally stimulated triggered 


443

444 

release have been attempted, including hyperthermic conditions 6 used for thermosensitive liposomes 

or polymeric micelles, and sonication 7 used to trigger release from Pluronic w micelles. 

In this chapter, polymeric micelles for the triggered release either by slightly acidic extracellular 

pH (pH e) or by early endosomal pH, are introduced. Triggered release by tumor pH e may allow 

localized high-dose therapy, which is effective for sensitive cells. Active intracellular translocation 

of the micellar carriers via specific interactions, combined with triggered release in early endosomes 

and endosomolytic activity of destabilized micelle components, has been proven to be 

effective for treating multidrug-resistant tumors. 

22.2 ACIDIC TUMOR pHe 

The pHe of normal tissues and blood is around pH 7.4, and the intracellular pH (pHi) of normal cells 

is around pH 7.2. However, in most tumor cells the pH gradient is reversed (pHiOpHe). Particularly, 

pHe is lower than in normal tissues. 8–11 Although there is a distribution of in vivo pHe measurements of human patients having various solid tumors (adenocarcinoma, squamous cell 

carcinoma, soft tissue sarcoma, and malignant melanoma) in readily accessible areas (limbs, 

neck, or chest wall) using needle type microelectrodes, the mean pH value is reported to be 7.0 

with a full range of 5.7–7.8, 11 the variation being dependent on tumor histology and tumor volume. 

Recent measurements of pHe by noninvasive technologies, such as 19 F, 31 P, or 1 Hprobesin 

magnetic resonance spectroscopy in human-tumor xenografts in animals, further verified consistently 

low pHe. 12,13 All measurements of pHe of human and animal solid-tumors by either invasive 

or noninvasive method showed that more than 80% of all measured values fall below pH 7.2. This 

tumor pH can be further lowered by hyperglycemic conditions. 14 

The high rate of glycolysis in tumor cells, under either aerobic or anaerobic conditions, has 

been thought to be a major cause of low pHe. The tumor cells synthesize ATP both by mitochondrial 

oxidative phosphorylation and glycolysis. Glycolysis produces two moles of lactic acid (pKaZ3.9) 

and two moles of ATP upon consumption of one mole of glucose. The hydrolysis of generated ATP 

produces protons. Despite the high production rate of protons in tumor cells, their cytosolic pH, 

particularly of resistant cells, is maintained to be alkaline, creating a favorable condition for 

glycolysis. The mechanisms responsible for exporting protons out of the cells remain unknown. 

This, along with the help of inadequate blood supply, poor lymphatic drainage, and high interstitial 

pressure in tumor tissues, leads to low pHe. 8–10 However, it is of interest to note that in one study, 

glycolysis-deficient variant cells (lacking lactate dehydrogenase) produced negligible quantities of 

lactic acid but still presented an acidic environment. 15 This finding has indicated the acidity may be 

a phenotype of tumor cells rather than a consequence of cell metabolic events. The acidic environment 

may benefit the cancer cell because it promotes invasiveness (metastasis) by destroying the 

extracellular matrix of surrounding normal tissues. 

A few attempts for pH-triggered release using pH-sensitive liposomes have been described in 

the literature, 16,17 but a practical system that triggers the release at tumor pHe is far off because of 

technical difficulties in endowing the desired pH sensitivity to the liposomes. 

22.3 MULTIDRUG RESISTANCE AND CANCER 


The resistance of cancer cells to multiple structurally unrelated chemotherapeutic drugs has been 

termed multidrug resistance (MDR). Clinical resistance of some cancers to cytotoxic drugs has 

been observed since the beginning of chemotherapy. The degree of resistance varies widely 

depending on the type of cancer. In past years, various mechanisms for MDR have been proposed, 

including the expression of cell surface multidrug efflux pumps (P-glycoprotein (Pgp) 18 and 

multidrug resistance-associated proteins (MRP1) 19 ) altered glutathione metabolism, 20 reduced 

activity of topoisomerase II, 21 and various changes in cellular proteins 22–24 and mechanisms. 25 


Polymeric Micelles Targeting Tumor pH 445 

Strong evidence shows that the expression of (Pgp) is associated with drug-resistance in cancer. 

Many attempts to circumvent Pgp-based MDR in cancer chemotherapy have utilized Pgp blocking 

agents (MDR modulators or reverters), 26–30 that can be co-administered with the anti-cancer drug. 

This approach is based on the premise that inhibiting Pgp function will result in increased accumulation 

of most anti-cancer drugs in the tumor cells and restore anti-tumor activity. However, 

co-administration of MDR modulators with anti-cancer drugs has often resulted in exacerbated 

toxicity of the anti-cancer drugs and very little improvement in chemosensitization of MDR cancers 

because of the blockade of Pgp excretory functions in healthy tissues, such as liver and kidney, 

which markedly reduces anti-cancer drug clearance. 

Independently, acidic intracellular organelles also participate in resistance to chemotherapeutic 

drugs. Parental drug sensitive cells are characterized to have rather acidic and diffused pH profile 

inside the cells, while multidrug resistant cells develop more acidic organelles (recycling endosome, 

lysosome, and trans-Golgi network) compared to cytosolic and nucleoplasmic pH. 31–33 

Because most anti-cancer drugs are in an ionizable form, the pH of extracellular matrix and 

intracellular compartments is a critical factor in determining drug partitioning and distribution. For 

instance, weakly-basic drugs accumulate more in an acidic phase, where the ionized form predominates, 

because an ionized drug is less permeable through cellular or subcellular membranes. This 

is more pronounced in MDR cells. 

It is claimed that the drug sequestration followed by transport to the plasma membrane and 

extrusion into the external medium is another major mechanism for MDR for weakly-acidic 

drugs. 31,34 It is interesting to note that transfection of P-gp into certain sensitive cells is often 

coupled with increased cytosolic pH. However, this is not the case for MCF-7 cell line. The 

transfection did not influence the internal pH but induced 20-fold increased resistance. When the 

same cell line was selected from an escalated doxorubicin treatment schedule, the cells presented 

the overexpression of P-gp together with typical MDR pH profile and increased the resistance 980fold. 

These observations indicate that the activity of the efflux pump, combined with drug sequestration 

and exocytosis mechanism, seem to be highly influential factors in MDR. 35 Thus, this 

subcellular pH effect must be taken into account in the carrier-design to avoid the sequestration 

effect and to improve drug availability to the cytosol and nucleus. 34 Few investigators have 

addressed the drug sequestration issue in drug-carrier design. 

This chapter summarizes our attempt for the triggered release of doxorubicin, either by tumor 

pHe or by early endosomal pH, from pH-sensitive polymeric micelles which are based on poly(Lhistidine) 

core. 

22.4 POLYHIS-PEG-BASED pH-SENSITIVE MICELLES: PREPARATION AND 

CHARACTERISTICS 

22.4.1 SYNTHESIS 

L-Histidine (His) was derivatized by introducing carbobenzoxy (CBZ) group to a amino group, and 

amino group in imidazole ring of N a -CBZ-L-histidine was protected with dinitrophenyl (DNP) 

group. N a -CBZ-N im -DNP-L-His was then transformed to the N-carboxy anhydride (NCA) form 

by thionyl chloride. The ring-opening polymerization of N a -CBZ-N im -DNP-L-His NCA$HCl in 

dimethylformamide (DMF) was performed using different molar ratios of the monomer to initiator 

(isopropylamine or n-hexylamine) (M/I ratio). 36 

Monocarboxy-PEG was used to prepare an activated N-hydroxysuccinyl (NHS)–poly(ethylene 

glycol) (PEG). The coupling reaction between poly(N im -DNP-L-histidine) and NHS–PEG (1:1 

functional group ratio) was carried out in tetrahydrofuran (THF) for two days at room temperature. 

After the reaction, the block copolymer was precipitated in a mixed solvent (THF/n-hexane) and 

filtered. 2-Mercaptoethanol was added to the poly(N im -DNP-L-histidine)-PEG diblock copolymer 

in DMF to deprotect the imidazole group (Scheme 22.1). Deprotection was complete within one 


446 

day at room temperature. The polymer solution was added to diethyl ether (800 mL) at 08C to 

precipitate polymer, while excess 2-mercaptoethanol and DNP-mercaptoethanol remained soluble. 

The polymer was filtered, washed with diethyl ether, and dried in vacuo for two days. For further 

purification, the polymer was dissolved in a minimum volume of 3 N HCl and stored in a 08C for 

one day. After the precipitation of DNP was complete, the solution was filtered through a 0.45-mm 

filter and the product was precipitated by adding an excess amount of acetone to the solution, and 

then dried in vacuo for two days. Dialysis (Spectra/Por; molecular weight cut off (MWCO) 5000) 

was used to remove uncoupled polymers. Before subsequent experiments, the HCl salt of polyhistidine 

(1 mmol) in dimethylsulfoxide (DMSO) was converted to the free base by treatment 

with TEA (2 mmol) for 3 h at room temperature; the free base was recrystallized from a solution 

mixture of ethanol and DMSO (10:1). 

It has been documented in literature that histidine residues in proteins significantly contribute to 

the protein buffering capacity 37 at physiological pH. The acid-base titration profiles of the polyHis 

homopolymer and polyHis/PEG block copolymers are presented in Figure 22.1. All of the polymer 

solutions exhibited a buffering pH region of pH 4–9. The titration curve confirmed that the polymers 

with a higher molecular-weight polyHis block had a higher buffering capacity (Figure 22.1a) 

in the physiological pH range of pH 5.5–8.0 due to the higher concentrations of imidazole group. 

PolyHis (M nw5000)/PEG (M nw2000) (PolyHis5K/PEG2K) and PolyHis (M nw3000)/PEG 

(M nw2000) (polyHis3K/PEG2K) had inflexion points around pH 7.0 (pK b). PolyHis5K and Poly- 

His3K showed pK b values around pH 6.5. The pK b shift of the block copolymer compared to the 

polyHis homopolymer may be due to increased hydration by PEG. 38 

22.4.2 POLYHIS/PEG COPOLYMER MICELLES 

The deprotonated polyHis at high pH is hydrophobic, whereas PEG is soluble in water at all pHs. 

This amphiphilicity is responsible for the formation of polymeric micelles. Lowering the solution 

pH below the pKb can affect the micellar structure because protonation decreases the hydrophobicity 

of polyhistidine. 36 

The polyHis/PEG block copolymer micelles were prepared by the dialysis of the copolymer 

solution in DMSO against a pH 8.0 medium. This was done in the presence of pyrene as 

H2N CH COOH 

CH2 carbobenzyl 

chloride 

N NH 

O 

O C C O 

HN 

CH2 N N 

initiator 

NO 2 

O 

H 

CH2 O C N CH COOH 

CH2 N NH 

N α -CBZ-L-histidine 

O 

H 

N CH C initiator residue (R) 

CH2 n 

N N 

NO 2 

NO2 NO2 N α -CBZ-N im -DNP-L-histidine NCA poly(N im -DNP-L-histidine) 

deblocking 

O 

PEG C 

H 

H 

O 

N CH C R 

CH2 n 

N NH 

PEG-b-poly(L-histidine) 

2,4-dinitrofluoro 

benzene 

carboxy group 

activated PEG 


O 

H 

CH2 O C N CH COOH 

CH2 N N 

NO 2 

NO2 N α -CBZ-N im -DNP-L-histidine 

O 

PEG C 

N N 

thionyl 

chloride 

O 

H 

N CH C R 

CH2 n 

NO 2 

NO 2 

O 

O C C O 

HN 

CH2 N N 

NO 2 

NO2 N im -DNP-L-histidine NCA 

SCHEME 22.1 Overall scheme for the syntheses of protected L-histidine monomer, protected poly(L-histidine) 

by ring-opening polymerization of NCA, coupling with PEG, and imidazole ring deprotection. (From 

Lee, E. S., Shin, H. J., Na, K., and Bae, Y. H., J. Control Release, 90, 363, 2003. With permission.) 



pH 

14 

12 

10 

8 

6 

4 

2 

0 

0 200 400 600 800 

0 

0 200 400 600 

(a) 1 N HCL (μL) 

(b) 

1 N HCL (μL) 

FIGURE 22.1 The pH profile of (a) polyHis5K/PEG2K (C), polyHis5K (&), NaCl (;) and (b) polyHis3K/ 

PEG2K (C), polyHis3K (&) by acid-base titration. (From Lee, E. S., Shin, H. J., Na, K., and Bae, Y. H., 

J. Control Release, 90, 363, 2003. With permission.) 

a fluorescent probe so that the micelle formation could be monitored by fluorometry. Pyrene 

strongly fluoresces in a nonpolar environment, while in a polar environment it shows weak fluorescence 

intensity. The change of total emission intensity vs. polymer concentration indicates the 

formation of micelles or the change from micelles to unimers (the dissociatation of polymers from 

the micelles). 39 PolyHis3K/PEG2K micelles exhibited instability over time, as evidenced by 

increasing light transmittance, while the stability of polyHis5K/PEG2K micelles was maintained 

for two days (data not shown). The instability of polyHis3K/PEG2K was probably due to short 

polyHis block length and indicates that there may be a critical polyHis length for stability 

somewhere between Mn 3000 and 5000. Therefore, polyHis5K/PEG2K was used for 

further investigation. 

The effect of the dialysis pH on critical micelle concentration (CMC) was examined and the 

results are presented in Figure 22.2. At a dialysis pH between 8.0 and 7.4, the CMC of polyHis5K/ 

PEG2K micelle increased slightly with decreasing pH. However, the CMC was significantly 

elevated below pH 7.2. It is evident that the protonation of the imidazole group in the copolymer 

at lower pH causes a reduction in hydrophobicity, leading to an increase in CMC. 

CMC (μg/mL) 

150 

100 

50 

0 

5.0 5.5 6.0 

pH 

pH 

14 

12 

10 

8 

6 

4 

2 

6.5 7.0 7.5 8.0 

FIGURE 22.2 The pH effect on the CMC of polyHis5K/PEG2K polymeric micelles that were prepared at 

various pHs (pH 8.0–5.0). (From Lee, E. S., Shin, H. J., Na, K., and Bae, Y. H., J. Control Release, 90, 363, 

2003. With permission.) 


800

448 

Intensity (%) 

(a) 

25 

20 

15 

10 

5 

0 

1 

10 100 

Particle size (nm) 

1000 

In addition, below pH 5.0 (typically pH 4.8), the CMC of polyHis5K/PEG2K micelle could not 

be detected. Taken together, at pH 8.0 the less protonated polyHis constitutes the hydrophobic core 

in the micellar structure, but in the range of pH 5.0–7.4 the polymer produced less-stable micelles. 

The size of the micelles formed from polyHis5K/PEG2K measured by a zetasizer was around 

110 nm, based on the intensity-average diameter with a unimodal distribution (Figure 22.3a), and 

the ratio of weight-average particle size to number-average particle size was 1.2. The relatively 

large size of polyHis5K/PEG2K micelles may be the result of dialysis conditions, which probably 

produced a secondary micelle structure that is a cluster of individual single micelles. 40 However, 

the polyHis5K/PEG2K micellar shape was regular and spherical, as visualized in the atomic force 

microscopy photograph (Figure 22.3b). The morphology of polymeric micelles was further evident 

from dynamic light scattering (DLS) measurements. 

A noticeable increase in light-scattering intensity was observed after dialyzing the block copolymer 

solutions at pH 8.0, while the unimer solution in DMSO was practically transparent. The 

pH-dependent micelle stability was confirmed again by the measurement of transmittance of the 

micelle solution. Figure 22.4a shows the transmittance change of the polymeric micelle solution as 

a function of pH. As the pH of the micelle solution decreased from the dialysis conditions (pH 8, 

ionic strength 1.0), the transmittance increased. In particular, this increase was prominent from pH 

7.4 to 6.0, reaching a plateau around pH 6.4. This result implies that polyHis5K/PEG2K micelles 

start dissociation starting at pH 7.4 and show a sharp transition between pH 7.4 and 7.0 followed by a 

gradual transition between pH 7.0 and 6.0. 

Figure 22.4b shows the changes in the particle size distribution of the micelles as a function of 

pH. The instability of the micelles on reduction of pH resulted in release of polyHis5K/PEG2K 

unimers or unimicelles isolated from secondary micelles, resulting in a bimodal size distribution 

with existing secondary micelles. The intensity of the peak at size 2–30 nm gradually increased 

with micelle dissociation at a lower pH. Below pH 7.4, an abrupt increase in the intensity of smallsized 

particles was apparent, which is consistent with the transmittance change of the micelle 

solution at this pH (Figure 22.4a). 

22.4.3 MIXED MICELLES OF POLYHIS/PEG AND PLLA/PEG 

The PolyHis/PEG micelle was relatively unstable and began to gradually disintegrate at pH 7.4. 

The incorporation of a nonionizable block copolymer (Poly(L-lactic acide) (PLLA)/PEG) into the 

Å 

20.0 

10.0 

0.0 

(b) 


0 0.2 0.4 µm 

FIGURE 22.3 The particle size distribution of polyHis5K/PEG2K polymeric micelles measured by (a) zetasizer 

and (b) AFM. (From Lee, E. S., Shin, H. J., Na, K., and Bae, Y. H., J. Control Release, 90, 363, 2003. 

With permission.) 



Transmittance % 

102 

100 

98 

96 

94 

92 

0 

6.0 6.5 7.0 7.5 8.0 1 

10 

100 1,000 

(a) pH 

(b) 

Particle size (nm) 

FIGURE 22.4 pH-dependent micelle destabilization. (a) The transmittance change of polyHis5K/PEG2K 

polymeric micelles (0.1 g/L) with pH. The polyHis5K/PEG2K polymeric micelles were prepared in 

NaOH–Na2B4O7 buffer solution (pH 8.0) and exposed at each pH for 24 h. (b) The change of particle size 

distribution in polyHis5K/PEG2K micelles exposed to each pH solution for 24 h; pH 8.0 (C), 7.4 (&), 7.2 

(:), 6.8 (;), 6.4 (%) and 6.0 (B). (From Lee, E. S., Shin, H. J., Na, K., and Bae, Y. H., J. Control Release, 

90, 363, 2003. With permission.) 

micellar structure of polyHis/PEG improved micelle stability at pH 7.4 and shifted destabilization 

to lower pHs. In particular, the mixed micelles composed of polyHis/PEG (75 wt%) and 

PLLA/PEG (25 wt%) showed a unimodal size distribution with an average size of 70-nm diameter 

at pH 9.0. Other mixed micelles prepared from different contents of PLLA/PEG ranged 70–100 nm 

in diameter with unimodal size distributions at pH 9.0. 41 

Figure 22.5 shows the effect of PLLA/PEG content on the CMC of the mixed micelles. When 

the mixed micelle was prepared at pH 9.0 and tested at pH 9.0, the CMC of the mixed micelles was 

very slightly influenced by PLLA/PEG content in a range of 0–25 wt%. However, above 25 wt% 

PLLA/PEG the CMC gradually increased, reaching to the higher CMC value of the PLLA/PEG 

micelle, although the CMC change is not statistically significant (Figure 22.5a). This indicates that 

a small amount of PLLA/PEG does not disturb the core stability. 

The presence of PLLA/PEG did significantly influence the pH-dependent CMC of the micelles. 

The CMC of polyHis/PEG micelle sharply increased from 2.7 to 140 mg/mL as pH decreased from 

8 to 5, with a major transition occurring between pH 7.0 and 6.8 (Figure 22.5b). However, the mixed 

micelle having 25 wt% PLLA/PEG showed a small increase in CMC, from 2.6 to 3.5 mg/mL, and 

the transition occurred in a pH range of 7.5–6.8 (Figure 22.5c). 

Figure 22.6 shows the pH-dependent stability of the mixed micelles monitored by relative light 

transmittance (T/T i%; the transmittance of the mixed micelle at each pH/transmittance at pH 9.0). 

The T/Ti% of polyHis/PEG micelle slightly increased with decreasing pH, especially below pH 7.6, 

which is attributed to micelle destabilization followed by dissociation. PLLA/PEG micelle, being 

pH insensitive, showed no change in turbidity and preserved its stability in the entire pH 

range tested. 

It is interesting to note that the T/Ti% of the mixed micelles decreased rather than increased with 

decreasing pH. This result seems to be associated with ionization of polyHis/PEG, accompanied by 

separation and isolation of PLLA/PEG from the micelles, which in turn became segregated in 

water, and indirectly reflects the disintegration of the micelles by ionization of the imidazole 

groups along polyHis chains. The destabilizing pH was influenced by the amount of PLLA/PEG 

block copolymer in the mixed micelles. The addition of up to 10 wt% of PLLA/PEG to the micelle 

showed a slight shift in destabilizing pH. Although a 25 wt% addition considerably enhanced the 

micelle stability at pH 7.4 and destabilization only occurred below pH 7.0. At 40 wt% PLLA/PEG, 


Intensity (%) 

40 

30 

20 

10

450 

CMC (μg/ml) 

4.0 

3.5 

3.0 

2.5 

2.0 

0 

20 

40 60 80 100 

(a) Adding content (wt%) of PLLA/PEG 

(b) 

T / Tinf % 

CMC (μg/mL) 

(c) 

110 

100 

90 

80 

3.8 

3.6 

3.4 

3.2 

3.0 

2.8 

2.6 

5.0 

5.5 6.0 

6.5 7.0 7.5 8.0 

5 6 7 8 9 

pH 

FIGURE 22.6 pH-dependent relative turbidity (T/Ti%) of the mixed micelles composed of polyHis/PEG and 

PLLA/PEG (PLLA/PEG content: 0 wt% (B); 5 wt% (C); 10 wt% (&); 25 wt% (:); 40 wt% (;); and 

100 wt% (,)). T i is the transmittance at 500 nm at pH 9.0 and T is the transmittance at a given pH after 

micelles were stabilized for 24 h. (From Lee, E. S., Na, K., and Bae, Y. H., J. Control Release, 91, 103, 2003. 


CMC (μg/ml) 

pH 

160 

140 

120 

100 

80 

60 

40 

20 


0 

5.0 5.5 6.0 6.5 7.0 7.5 8.0 

FIGURE 22.5 The CMC change of the mixed micelles prepared with (a) different contents of PLLA/PEG in 

NaOH–Na2B4O7 pH 9.0 buffer, (b) pH-dependent CMC of polyHis/PEG micelle in two solutions (HCl or 

NaOH–Na 2B 4O 7 ionic strengthZ0.1), and (c) pH-dependent CMC of the mixed micelle composed of 25 wt% 

PLLA/PEG and 75 wt% polyHis/PEG in the same solutions (nZ3). (From Lee, E. S., Shin, H. J., Na, K., and 

Bae, Y. H., J. Control Release, 90, 363, 2003. With permission.) 


pH


the destabilization pH was shifted a bit further downward with less release of PLLA/PEG due to its 

high content. 

22.5 TUMOR pHe TARGETING 

The pH-dependent micelle property prompted us to test the micelles with pH-induced drug release. 

When doxorubicin (DOX) was incorporated into the mixed micelles (0, 10, 25, and 40 wt% 

PLLA/PEG) during dialysis, the micelle sizes ranged from 50 to 70 nm as determined by 

dynamic light scattering. The drug loading efficiency was about 75–85% and the drug content in 

the micelles was 15–17 wt%. 41 

Figure 22.7a shows the cumulative DOX release from polyHis/PEG micelle. The DOX release 

pattern followed nearly first-order kinetics and reached a plateau in 24 h and 30 wt% of 

loaded DOX was released for 24 h at pH 8.0. Decreasing pH accelerated the release rate. The 

pH-dependent release patterns of the mixed micelles are presented in Figure 22.7b by plotting 

cumulative amount of DOX for 24 h vs. release medium pH. The mixed micelles containing 

25 wt% PLLA/PEG showed a favorable pH-dependency such that 32 wt% of DOX was released 

at pH 7.0, 70 wt% of DOX at pH 6.8, and 82 wt% at pH 5.0. On the other hand, the mixed micelles 

containing 40 wt% of PLLA/PEG suppressed the release of DOX to 35 wt% at pH 6.8, and 64 wt% 

at pH 6.6. This indicates that the content of PLLA/PEG controlled pH-dependent release from the 

mixed micelles by destabilization; the transition pH coincided with the results shown in 

Figure 22.6. The results support the idea that the mixed micelles truly recognize the minute 

difference in pH and discriminate the tumor pH by destabilization and release of micelle content. 

When the blank micelles were tested with human breast adenocarcinoma (MCF-7) cells, no 

noticeable cytotoxicity was observed up to 100 mg/mL of polymeric micelle for a 48-h culture, 

regardless of the culture medium pH. However, DOX-loaded micelles did present tumor cell 

toxicity in a pH-dependent manner. 

DOX-loaded polyHis/PEG micelles demonstrated a certain degree of cytotoxic effect at pH 7.4 

because of low micelle stability and DOX release. But the cell viability was much reduced at pH 6.8 

due to enhanced DOX release (Figure 22.8). The mixed micelles with 10 wt% PLLA/PEG 

(Figure 22.8b) caused low cytotoxicity at pH 7.4, which increases dramatically below pH 7.2. 

The above results are distinguished from the cytotoxicity of free DOX, which is almost independent 

of pH. The mixed micelles having 25 wt% PLLA/PEG (Figure 22.8c) appeared more sensitive to 

Cumulative DOX release (%) 

100 

80 

60 

40 

20 

Total DOX release (%) in 24 h 

0 

0 10 20 30 40 50 

20 

5.0 5.5 6.0 6.5 7.0 7.5 8.0 

(a) Time (h) 

(b) 

pH 

FIGURE 22.7 pH-dependent cumulative DOX release from the mixed micelles composed of polyHis/PEG 

and PLLA/PEG with varying PLLA/PEG content in the mixed micelles: (a) 0 wt%; pH 8.0 (C); 7.4 (&), 7.2 

(:), 7.0 (;), 6.8 (%), 6.2 (,), 5.0 (6). (b) 0 wt% (†), 10 wt% (&), 25 wt% (:) and 40 wt% (;) after 24 h. 

(From Lee, E. S, Na, K., and Bae, Y. H., J. Control Release, 91, 103, 2003. With permission.) 


100 

80 

60 

40

452 

Cell viability (%) by MTT assay 

(a) 


(e) 

100 

80 

60 

40 

20 

1 

20 

1 

10 100 1,000 10,000 

DOX ng/mL 

(c) DOX ng/mL 

(d) 


100 

80 

60 

40 

100 

80 

60 

40 

10 100 1,000 10,000 

20 

1 10 100 

DOX ng/mL 

1,000 10,000 

pH, differentiating especially between pH 7.0 and 6.8. The mixed micelles with 40 wt% PLLA/PEG 

micelles (Figure 22.8d) distinguished between pH 6.8 and 6.6 in a similar cytotoxicity study. 

Considering that a PLLA/PEG micelle (Figure 22.8e) has no pH-dependent cytotoxicity and 

neither do blank micelles, the pH-dependent cytotoxicity solely relies on the release of DOX at 

different pHs. The change in pH from 8.0 to 6.6 may influence cell surface charges, and cellular 

physiology and viability, noting that the low pH is a favorable environment for tumor cells but 

opposite for normal cells. The cell viability expressed in this study is relative to those at different 


(b) 


100 

80 

60 

40 

20 

1 10 100 1,000 10,000 

100 

80 

60 

40 


DOX ng/mL 

20 

1 10 100 1,000 10,000 

DOX ng/mL 

FIGURE 22.8 The cytotoxicity of DOX-loaded mixed micelles with different PLLA/PEG contents (a) 0 wt%, 

(b) 10 wt%, (c) 25 wt%, (d) 40 wt%, and (e) 100 wt% after 48 h incubation at varying pH (pH 7.4 (C); 

7.2 (&); 7.0 (:); 6.8 (;); and 6.6 (%)). Free-DOX cytotoxicity is presented as a reference at pH 7.4 (B), 

7.2 (,), 7.0 (6), 6.8 (7), 6.6 (&). (From Lee, E. S., Na, K., and Bae, Y. H., J. Control Release, 91, 103, 2003. 




pHs in the absence of DOX, and the direct pH effect on the cell viability was not monitored with the 

MCF-7 cell line. No noticeable difference in cell viability by pH with free DOX was observed. 

In vitro results of the mixed micelles with 25 wt% PLLA/PEG (referred as PH-sensitive mixed 

micelle (PHSM) hereafter) showed the pH-dependent cytotoxicity of DOX against sensitive MCF-7 

cells, and were comparable with free DOX and a conventional pH-insensitive PLLA/PEG micelle 

(referred to as pH-insensitive micelle (PHIM). The free DOX showed the highest destruction, which 

is not significantly influenced by pH variation between 6.8 and 7.4. PHIM showed the least cytotoxicity 

at the two tested pH values. However, PHSM showed a distinct pH effect by switching the 

cytotoxicity from the level of PHIM at pH 7.0 (MCF-7 cell viability was 85% after the treatment of 

equivalent DOX 1 mg/mL) to a similar level of free DOX at pH 6.8 (MCF-7 cell viability was 36% 

after the treatment of equivalent DOX 1 mg/mL). This was attributed to pH-triggered release of DOX. 

Figure 22.9a shows the in vivo results of the anti-cancer activity of the tested DOX formulations 

i.v. injected on days 0 and 3. The DOX loaded PHSM (equivalent DOXZ10 mg/kg) 

exhibited significant inhibition (P!0.05 compared with free DOX or saline solution) of the 

growth of s.c. MCF-7 xenografts. The tumor volume of mice treated with the PHSM (P!0.05 

compared with free DOX) was approximately 4.5 and 3.6 times smaller than that of saline solution 

or free DOX treatment after 6 weeks, respectively. The triggered release of DOX by tumor pHe (! 

7.0), after accumulation of the micelle in the tumor sites via the enhanced permeation and retention 

(EPR) mechanism, 42,43 may present a more effective modality in tumor chemotherapy, providing 

higher local concentrations of the drug at tumor sites (targeted high-dose cancer therapy), while 

providing a minimal release of the drug from micelles while circulating in the blood (pH 7.4). When 

compared with PHIM, the volume of tumors treated with the PHSM was approximately 3.0 times 

smaller (particle sizeZ50–70 nm, data now shown) than those whose properties are not influenced 

by pH. In addition, the destabilization of PHSM (converting to unimers) may help extravasation of 

additional micelles from the blood compartment and their penetration into tumors by reducing the 

physical barrier effect, which can be generated by stable particles, such as conventional liposomes 

and polymeric micelles. It was reported that drug carriers, such as PEG modified liposomes, are in 

an insoluble form and only reside in the vicinity of the leaky blood vessels after extravasation, 

rather than migrating into the deeper sites of tumors. This “road block effect” could obstruct the 

subsequent accumulation of the carriers at tumor sites. 44 It is interesting to note that only PHSM 

administration showed decreased tumor size for the initial first week. No obvious changes in mice 

Tumor volume (mm 3 ) 

1200 

1000 

800 

600 

400 

200 

0 

14 

0 5 10 15 20 25 30 35 40 45 0 10 20 30 40 


(b) 

Time (day) 

FIGURE 22.9 Tumor growth inhibition of s.c. human breast MCF-7 carcinoma xenografts in BALB/c nude 

mice. Mice were injected i.v. with 10 mg/kg DOX equivalent dose: PHSM (C), PHIM (&), free DOX (:), 

and saline solution (;). Two i.v. injections on day 0 and 3 were administered. (a) Tumor volume change and 

(b) body weight change (nZ5). Values are the meansGthe standard deviations (SD). (From Lee, E. S., Na, K., 

and Bae, Y. H., J. Control Release, 103, 405, 2005.) 


Body weight (g) 

26 

24 

22 

20 

18 

16

454 

body weight in experimental groups, except for the free DOX group, were observed. Free DOX 

caused more weight-loss after i.v. administration (Figure 22.9b) than micelle formulations. 

“Stealth” particles decorated with PEG are known to be invisible to macropharges and have 

prolonged half-lives in the blood compartment. This property endows the particles a higher probability 

to extravasate in pathological sites, like tumors or inflamed regions where a leaky vascular 

structure persists. 45 Recent investigations have reported that the size cut-off of tumor vasculature for 

the EPR effect grew to 200–700 nm. 46 Nevertheless, the nanoparticle penetration to tumor vasculature 

is heterogeneous in a solid tumor, and the vascular structure or cut off size differs from tumor 

to tumor and is influenced by tumor location. 46 A more recent study demonstrated that the optimal 

size of stealth liposomes was determined to be around 100 nm in an animal model. 47 This means that 

the size of micelle or liposome affects the in vivo bio-distribution, post injection. Smaller size causes 

more accumulation in the liver, while larger ones cause more accumulation in the spleen. 48 

We hypothesized that the average particle size of 50–70 nm of PHSM may meet the optimal 

size for extravasation or long-circulation suggested by several research groups. 42,44,47,48 A significant 

portion of DOX carried by the PHSM, however, was accumulated in the liver and spleen. In 

Figure 22.10a, DOX levels in the blood after i.v. administration of PHSM to the animals bearing s.c. 

MCF-7 carcinoma xenografts were 27% at 4 h, 8% at 12 h, 4% at 48 h in the blood and 18% at 4 h, 

15% at 12 h, 3% at 48 h in the tumors and 14% at 4 h, 7% at 12 h, 1.5% at 48 h in the liver and 2% at 

4 h, 1% at 12 h, barely detectable levels at 48 h in the spleen. The mononuclear phagocytes system 

(MPS), especially in the liver and spleen, 42,44,46,49 seems to be responsible for the accumulation in 

these organs even with small size of the particles. 

As shown in Figure 22.8b, because of very short half-life (t1/2, within few minutes) of free DOX 

in the blood, 50 DOX levels in the blood at 12–48 h were not detectable. DOX levels in the tumors 

were 1.5% at 4 h, 1% at 12 h, and barely detectable levels were found at 48 h. In contrast, high 

concentrations of DOX were found in the heart, liver and considerable DOX levels were shown in 

Total DOX (%) in orans 

50 

40 

30 

20 

10 

(a) 0 

40 

30 

20 

10 

(b) 0 

tumor blood liver heart spleen lung kidney 


FIGURE 22.10 Total DOX (%) levels (total DOX remaining in organs to DOX amount first injected) in blood 

and tissues (tumor, liver, heart, spleen, lung, kidney) after i.v. administration (10 mg/kg DOX equivalent dose) 

of (a) PHSM micelles, and (b) free DOX. The mice (s.c. human breast MCF-7 carcinoma xenografts) were 

sacrificed at 4 (&), 12 (&), and 48 h (&) after i.v. administration. Values are the meansGthe standard 

deviations (nZ4, SD). (From Lee, E. S., Na, K., and Bae, Y. H., J. Control Release, 103, 405, 2005.) 



the lung, spleen and kidney. These results are in striking contrast to DOX biodistribution when 

carried by PHSM. Especially, PHSM is responsible for the enhanced tumor accumulation. 51 The 

combined mechanisms of pH e-triggered release and EPR effect can be beneficial to treat solid 

tumors by minimizing the road block effect of the particles after extravasation, increasing local 

concentration for diffusion to the tumors and more active internalization. This, however, requires 

further investigation for verification. 

22.6 EARLY ENDOSOMAL pH TARGETING FOR REVERSAL OF MDR 

PolyHis has been known for endosomolytic activity and recently low MW polyHis-conjugated 

poly(L-lysine) has demonstrated improved gene transfection. 52 However, the difficulty in controlling 

its molecular weight and limited solubility of high MW polymer in organic solvents has restricted the 

fabrication and applications of polyHis-based drug carriers. The mixed micelles, who’s surface was 

decorated with folate, consisting of polyHis/PEG-folate (Figure 22.11a) and PLLA/PEG-folate 

(Figure 22.11b), showed a similar release pattern of the micelles without folate (PHSM); cumulative 

DOX release for 24 h of 30–35 wt% in the pH range of 7.0–7.4, which is slightly more than 20– 

30 wt% initial burst release from conventional pH-insensitive micelles, 53 but about 70 wt% at pH 

6.8 during the same time period. During this process, it was observed that the drug or polymer was 

distributed in the cytosol and the nucleus in a sensitive breast cancer cell line (MCF-7). 41 

From these results, we hypothesized that the folate receptor-mediated internalization route and 

the micelle destabilization reverse the drug resistance of solid tumors. For proof of concept, a MCF- 

7/DOX R cell was selected after being stepwise exposed to 0.001–10 mg/mL of free DOX as 

reported. 54–58 

Figure 22.12 shows the decreased cytotoxicity of free DOX against MCF-7/DOX R cells, exhibiting 

a typical drug resistance curve. The cytotoxicity of DOX against sensitive MCF-7 cells 

slightly decreased by lowering pH from 8.0 to 6.8 because of an increased degree of ionization 

of DOX, causing less diffusivity through the plasma membrane. 59 For the resistant cells, the pH 

effect was not noticed even at a high DOX concentration range. The IC50 value of free DOX 

increased from 470 ng/mL for wild MCF-7 cells to 25,800 ng/mL for MCF-7/DOX R cells at pH 7.4. 

Figure 22.13 visualizes different distributions of free DOX in MCF-7 (Figure 22.13a) and 

MCF-7/DOX R (Figure 22.13b) cells. In the MCF-7/DOX R cells, there was a limited DOX uptake 

in the cytosolic compartment, while the sensitive MCF-7 cells accumulated DOX in the cytosolic 

compartment and nucleus. Figure 22.14 shows the cytotoxic effects of free DOX or DOX-loaded 

(a) 

(b) H2N HN 

H 2 N 

HN 

O 

N 

O 

N 

N 

N 

N N 

O 

N 

H N 

H 

COOH 

O 

O COOH 

N 

H N 

H 

O 

O 

PEG C 

H 

O 

N H C C 

CH 2 

N N H 

FIGURE 22.11 The chemical structures of (a) polyHis/PEG-folate and (b) PLLA/PEG-folate. 


O 

H 

N 

N 

H 

O 

C 

PEG 

O 

O 

H CH3 N CH 

n 

CH 3 

CH 3 

O 

n 

H

456 

Cell Viability (%) by MTT assay 

120 

100 

80 

60 

40 

20 

0 

1 10 100 1000 10000 100000 

DOX ng/mL 

FIGURE 22.12 Characterization of micelles on MCF-7/DOX R cells: The cytotoxicity of free DOX against 

MCF-7 (close circle) and MCF-7/DOX R (open circle) cells at pH 8.0 (C), pH 7.4 (&) and pH 6.8 (:)(nZ7). 

(From Lee, E. S., Na, K., and Bae, Y. H., J. Control Release, 103, 405, 2005.) 

FIGURE 22.13 Confocal microscopy studies on (a) DOX-sensitive MCF-7 and (b) DOX-resistant MCF-7 

(MCF-7/DOX R ) cells treated with free DOX at pH 6.8. 


(a) 

120 

100 

80 

60 

40 

20 

0 0 10 100 1000 10000 

DOX ng/mL 


(b) 

120 

100 

80 

60 

40 

20 


0 

1 10 100 1000 10000 

DOX ng/mL 

FIGURE 22.14 The cytotoxicity of DOX-loaded PHSM/f (C), polyHis/PEG-folate micelles (&), 

PLLA/PEG-folate micelles (:), free DOX (;) against MCF-7/DOX R cells at (a) pH 7.4 and (b) pH 6.8 

after 48 h incubation (nZ7). (From Lee, E. S., Na, K., and Bae, Y. H., J. Control Release, 103, 405, 2005.) 



micelles against MCF-7/DOX R cells at pH 7.4 and 6.8. The viability of MCF-7/DOX R cells was 

not dramatically decreased as expected at even high concentrations of free DOX due to the 

influence of Pgp function. DOX-loaded PLLA/PEG-folate micelles (pH-insentive micelles as a 

control) showed limited cytotoxicity against MCF-7/DOX R cells. DOX loaded polyHis/PEGfolate 

micelles demonstrated a similar degree of cytotoxicity against MDR cells as that of free 

DOX against the sensitive cells. 41 This effect is most likely due to enhanced DOX release and the 

endosomal escaping activity of polyHis/PEG-folate unimer released after internalization of the 

micelles and subsequent destabilization. This can be regarded as cytosolic delivery of DOX, 

disturbing drug sequestration mechanism of MDR. 41 However, it is shown that the viability of 

cells treated by PHSM with folate (PHSM/f) was about 10% at 10 mg/mL (DOX concentration), 

whereas for polyHis/PEG-folate micelles (without PLLA/PEG-folate) the value is about 40% at 

similar DOX concentration (Figure 22.14a). This enhanced effect suggests that there could be an 

additional mechanism of PHSM/f beside the functions of polyHis. Hoffman’s group reported that 

the membrane disruption mechanisms of poly(2-ethylacrylic acid) (PEAA) is presumably related 

to their increase in hydrophobicity at low pH values. 60 Similar to that, PLLA/PEG-folate released 

from destabilized PHSM/f could interact with the endosomal membrane due to the exposed 

hydrophobic PLLA portion, disturbing membrane stability. There is no significant difference 

between two pHs tested (pH 7.4 (Figure 22.14a) and 6.8 (Figure 22.14b)). Figure 22.14b suggests 

the feasibility of MDR reversal by PHSM/f at acidic tumor pH e. 61 

In the case of sensitive MCF-7 cells, PHSM/f showed significant cytotoxicity against cells. 

Figure 22.15 demonstrates this effect. The DOX-loaded micelles with folate were internalized via 

folate-receptor mediated endocytosis. The PHSM/f micelles induced greatly enhanced cytotoxicity 

(cell viability 16% at DOX 10 mg/mL at pH 8.0 (this pH was employed to ensure the micelle 

stability before internalization)) resulting from increased micelle concentration inside the cell due 

to folate receptor-mediated endocytosis and fusogenic effect of polyHis, while the PLLA/PEGfolate 

showed a slight increase in cytotoxicity. These observations clarify the role of polyHis in 

anti-cancer activity as two folds: (1) destabilization for drug release and (2) fusogenicity. 

Figure 22.15b indicates that, once internalization by folate receptors grew predominant, the pH 

effect on drug release became insignificant. This observation can be explained by the facts that (1) 

the mixed micelles were internalized into tumor cells via folate-receptor mediated endocytosis and 

underwent destabilization in a slightly acidic early endosomal compartment by protonation of an 

imidazole group and (2) DOX released from the mixed micelles under the influence of the extracellular 

pH, regardless of folate-receptor mediated endocytosis, has effective an drug portion for 

cell killing. Figure 22.15c shows the cell killing rates of the mixed micelles and free DOX. The free 

DOX diffuses only passively to cells. Unlike free DOX, PHSM/f facilitated faster cell killing in 

0–4 h of incubation. This observation is consistent with the high number of folate-receptors on 

tumor cells. In addition, there is little difference in the cell killing rate between pH 6.8 and 7.4 (data 

not shown), thus explaining quick cytotoxic action of DOX-loaded mixed micelles against MCF-7 

cells due to the active internalization. 58,59 

For further evidence of the functioning of PHSM/f, a confocal-microscopic study was 

conducted to reveal the distribution of DOX in the cells. Figure 22.16a shows the localized 

DOX in MCF-7/DOX R cells treated with DOX-loaded PLLA/PEG-folate micelles. The PLLA/ 

PEG-folate micelles were entrapped in endosome or multivesicular bodies resulting from folate 

receptor-mediated endocytosis. This could be responsible for the discrete DOX-intensity and biased 

DOX-intensity distribution. The entrapment seems to be a consequence of the absence of fusogenic 

activity. Unlike PLLA/PEG-folate micelles, PHSM/f incubation resulted in broad distribution of 

DOX in cytosol as well as nuclear compartments (Figure 22.16b). 

To bypass one of the major MDR mechanisms, Pgp pumping, we employed folate receptormediated 

endocytosis, because MCF-7 cells over-express folate receptors. 62 Along with this entry 

mechanism, the further enhanced drug release at endosomal pH 36,41 and the disruption of the 

endosomal membrane 41,51,52,63 are expected to kill MDR cells. 


458 


100 

80 

60 

40 

20 

0 

1 10 100 1000 10000 

(a) DOX ng/mL 

(b) 


(c) 

110 

100 

90 

80 

70 

60 

50 

40 

0 

1 10 

0 5 10 15 20 25 

Incubation time (h) 

FIGURE 22.16 Confocal microscopy studies on MCF-7/DOX R cells treated with (a) DOX-loaded 

PLLA/PEG-folate micelles and (b) DOX-loaded PHSM/f micelles at pH 6.8. The cells were incubated with 

the carriers for 30 min. (From Lee, E. S., Na, K., and Bae, Y. H., J. Control Release, 103, 405, 2005.) 


100 

80 

60 

40 

20 


100 

DOX ng/mL 

1000 10000 

FIGURE 22.15 The cytotoxicity of (a) free DOX (C), polyHis/PEG-folate micelle (&), PLLA/PEG-folate 

micelle (:) and PHSM/f (;) at pH 8.0 and of (b) PHSM/f at pH 7.4 (C) and pH 6.8 (&) against sensitive 

MCF-7 cells after 48 h incubation. (c) The cytotoxicity against MCF-7 cells treated with PHSM/f (C) and free 

DOX (&) at pH 6.8. The DOX content in the micelles was adjusted to be equivalent to free DOX concentration 

(500 ng/mL) and MCF-7 cells were treated for a given incubation time. (From Lee, E. S., Na, K., and Bae, Y. 

H., J. Control Release, 91, 103, 2003. With permission.) 



In the in vitro studies with MCF-7/DOX R cells, 51 the PHSM/f showed a significantly enhanced 

efficacy in tumor cell cytotoxicity, most likely due to enhanced internalization (folate receptormediated 

endocytosis), and subsequent pH-induced micelle destabilization and endosomal escape. 

However, the PHSM or PHIM with folate (PHIM/f) showed low efficacy in killing MCF-7/DOX R 

cells, due to lack of entry mechanism and/or endosomal escape. 

The MCF-7/DOX R xenografts in nude mice were used for the investigation of in vivo efficacy. 

The tumor bearing animals were treated by multiple i.v. injections on days 0, 3, and 6 

(Figure 22.17a). Neither complete tumor regression, nor toxicity-induced death was observed in 

all experimental groups, though, the tumor growth in mice treated by PHSM/f was inhibited and the 

size was reduced for two weeks after the third i.v. injection. The tumor volume in mice treated by 

PHSM/f (P!0.05 compared with free DOX) was approximately 2.7-times smaller than those 

treated with free DOX or PHIM, and approximately 1.9-times smaller than those treated with 

PHSM after six weeks. These results indicate that the folate-mediated endocytosis pathway of 

PHSM/f enhances the regression of tumors and is more efficient for MDR tumor chemotherapy 

than PHSM. This being said, the tumor regression efficacy in PHIM/f was not significant because 

some folate conjugates may have recycled back to the surface 64 before drug unloading. In addition, 

slow or limited drug release from PHIM maybe associated with drug sequestration in acidic 

organelles in MDR cells. 31,33 The active internalization, enhanced drug release rate, and endosomal 

escaping activity justify the efficacy of PHSM/f. Free DOX administration exhibited more weightloss 

in mice than the groups treated by any micelles (Figure 22.17b), signifying lower toxicity of 

DOX-loaded micelles. 

As shown in Figure 22.18, s.c. MCF-7/DOX R carcinoma xenografts showed different DOX 

levels in tumor as compared to s.c. MCF-7 carcinoma xenografts, particularly for PHSM. DOX 

levels in the tumors of s.c. MCF-7/DOX R carcinoma xenografts in which the DOX levels were 21% 

at 4 h, 12% at 12 h, and 5% at 48 h for PHSM/f; 7% at 4 h, 5% at 12 h, and 2% at 48 h for PHSM; 

and 1% at 4 h and barely detectable levels at 12 and 48 h for free DOX. This suggests that the role of 

the Pgp pump in s.c. MCF-7/DOX R carcinoma xenografts is significant in determining DOX levels 

in the resistant tumors. In the case of DOX delivered by PHSM into the resistant tumor, the released 

DOX cannot diffuse in the cells but is instead being washed away from the site. On the other hand, 

DOX carried by PHSM/f shows similar accumulation levels in the resistant tumor to that of 

sensitive tumor, resulting from active internalization by folate receptor-mediated endocytosis. 


1,000 

800 

600 

400 

200 

0 

0 5 10 15 20 25 30 35 40 45 

14 

0 10 20 30 40 


(b) 

Time (day) 

FIGURE 22.17 Tumor growth inhibition of s.c. human breast MCF-7/DOX R carcinoma xenografts in 

BALB/c nude mice. Mice were injected i.v. with 10 mg/kg DOX equivalent dose: PHSM/f (C), PHSM 

(&), PHIM/f (:), PHIM (;), and free DOX (%). Three i.v. injections on day 0, 3, and 6 were made. (a) 

Tumor volume change and (b) body-weight change (nZ5). Values are the meansGthe standard deviations 

(SD). (From Lee, E. S., Na, K., and Bae, Y. H., J. Control Release, 103, 405, 2005.) 


Body weight (g) 

26 

24 

22 

20 

18 

16

460 

Total DOX (%) in organs 

This, combined with the endosomal escaping capacity of PHSM/f, resulted in a remarkable 

regression effect for resistant tumors, as shown in Figure 22.18. 


40 

30 

20 

10 

(a) 0 

40 

30 

20 

10 

(b) 0 

40 

(c) 

30 

20 

10 

0 

Tumor blood liver heart spleen lung kidney 


FIGURE 22.18 Total DOX (%) levels (total DOX remaining in organs to DOX content first injected) in blood 

and tissues (tumor, liver, heart, spleen, lung, and kidney) after i.v. administration (10 mg/kg DOX equivalent 

dose) of (a) PHSM/f micelles, (b) PHSM micelles, and (c) free DOX. Mice (s.c. human breast MCF-7/DOX R 

carcinoma xenografts) were sacrificed at 4 (&), 12 (&), and 48 h (&) after i.v. administration. Values are the 

meansGthe standard deviations (nZ4, SD). (From Lee, E. S., Na, K., and Bae, Y. H., J. Control Release, 103, 

405, 2005.) 

Although important findings in scientific research and technological advances, such as long-circulating 

carries, EPR effect, tumor specific receptor-mediated targeting, and MDR modulator and 

reverters, have been achieved in the last few decades for effective solid tumor targeting, current 

chemotherapy is still facing major challenges to improving drug accumulation in tumor sites and 

reversing intrinsic or acquired drug-resistance of tumors. 

An accelerated release rate of doxorubicin at tumor extracellular pH was provided by pH-sensitive 

polyHis/PEG micelles, but the micelle was not fully stable at pH 7.4. When a pH-insensitive 

block copolymer, PLLA–PEG, was blended with polyHis-PEG to form a mixed micelle, the 

stability of the micelles at pH 7.4 was significantly improved and the micelles’ destabilization 

pH shifted from 7.2 to 6.6 with an increase in the amount of PLLA–PEG in the mixed micelle. 

The micelles were further modified with folate. These pH-sensitive micelles proved to be much 

more effective in regressing wild-type MCF-7 cells in vitro and in vivo studies with less weight-loss 

in animals, when compared with the pH-insensitive PLLA–PEG micelle. The folate effect was not 

apparent with the sensitive cells in the in vitro results. It seems that doxorubicin released either 

extracellularly or intracellularly is effective for cell-kill due to its high diffusivity through the 

plasma membrane. 

The pH-sensitive micelle-with-folate showed a dramatic effect when tested with drug-resistant 

tumors in vitro and in vivo, demonstrating a similar cell-kill rate as that of sensitive cells. 



The underlying mechanisms involve (1) fast internalization of the micelles to bypass the Pgp, (2) 

fast release rate of doxorubicin at early endosomes, and (3) quick disruption of the endosomal 

membrane by virtue of polyHis protonation and possible membrane interactions of freed polyHis- 

PEG. These mechanisms avoid the major MDR mechanisms of the Pgp pump and drug sequestration 

in acidic organelles and endosomal recycling, which resulted in uniform intracellular 

distributions of doxorubicin and polyHis-PEG, including the nucleus, In a resistant tumor xenograft 

model in mice, the tumor was regressed by iv administration of the folate-decorated pH-sensitive 

micelles with less weight-loss of the treated animal than control animals treated by pH-insensitive 

micelles and free doxorubicin. 

REFERENCES 

1. Levasseur, L. M., Slocum, H. K., Rustum, Y. M., and Greco, W. R., Modeling of the time-dependency 

of in vitro drug cytotoxicity and resistance, Cancer Res., 58, 5749, 1998. 

2. Gardner, S. N., A mechanistic, predictive model of dose–response curves for cell cycle phase-specific 

and -nonspecific drugs, Cancer Res., 60, 1417, 2000. 


Theory and practice, Pharmacol. Rev., 53, 283, 2001. 

4. Yokoyama, M., Novel passive targetable drug delivery with polymeric micelles, In Biorelated Polymers 

and Gels, Okano, T., Ed., Academic Press, San Diego, pp. 193–229, 1998. 

5. Drummond, D. C., Meyer, O., Hong, K., Kirpotin, D. B., and Papahadjppoulos, D., Optimizing 

liposomes for delivery of chemotherapeutic agents to solid tumors, Pharmacol. Rev., 51, 691, 1999. 


delivery: Effect of particle size, Cancer Res., 60, 4440, 2000. 

7. Munshi, N., Rapoport, N., and Pitt, W. G., Ultrasonic activated drug delivery from Pluronic P-105 

micelles, Cancer Lett., 118, 13, 1997. 

8. Hobbs, S. K., Monsky, W. L., Yuan, F., Roberts, W. G., Griffith, L., Torchilin, V. P., and Jain, R. K., 

Regulation of transport pathways in tumor vessels: Role of tumor type and microenvironment, Proc. 

Natl Acad. Sci. USA, 95, 4607, 1998. 

9. Tannock, I. F. and Rotin, D., Acid pH in tumors and its potential for therapeutic exploitation, Cancer 

Res., 49, 4373, 1989. 

10. Stubbs, M., McSheehy, R. M. J., Griffiths, J. R., and Bashford, L., Causes and consequences of tumour 

acidity and implications for treatment, Opinion, 6, 15, 2000. 

11. Engin, K., Leeper, D. B., Cater, J. R., Thistlethwaite, A. J., Tupchong, L., and McFarlane, J. D., 

Extracellular pH distribution in human tumors, Int. J. Hyperthermia, 11, 211, 1995. 

12. van Sluis, R., Bhujwalla, Z. M., Raghunand, N., Ballesteros, P., Alvarez, J., Cerdán, S., Galons, J. P., 

and Gillies, R. J., In vivo imaging of extracellular pH using 1 H MRSI, Magn. Reson. Med., 41, 743, 

1999. 

13. Ojugo, A. S. E., McSheehy, P. M. J., Mcintyre, D. J. O., McCoy, C., Stubbs, M., Leach, M. O., Judson, 

I. R., and Griffiths, J. R., Measurement of the extraceullar pH of solid tumours in mice by magnetic 

resonace spectroscopy: A comparison of exogenous 19 F and 31 P probes, NMR Biomed., 12, 495, 1999. 

14. Leeper, D. B., Engin, K., Thistlethwaite, A. J., Hitchon, H. D., Dover, J. D., Li, D. J., and Tupchong, 

L., Humon tumor extracellular pH as a function of blood glucose concentration, Int. J. Radiat. Oncol. 

Biol. Phys., 28, 935, 1994. 

15. Yamagata, M., Hasuda, K., Stamato, T., and Tannock, I. F., The contribution of lactic acid to 

acidification of tumours: Studies of variant cells lacking lactate dehydrogenase, Br. J. Cancer, 77, 

1726, 1998. 

16. Gerasimov, O. V., Boomer, J. A., Qualls, M. M., and Thompson, D. H., Cytosolic drug delivery using 

pH- and light-sensitive liposomes, Adv. Drug Deliv. Rev., 38, 317, 1999. 

17. Drummond, D. C., Zignani, M., and Leroux, J. C., Current status of pH-sensitive liposomes in drug 

delivery, Prog. Lipid Res., 39, 409, 2000. 

18. Gottesman, M. M., Pastan, I., and Ambudkar, S. V., P-glycoprotein and multi drug resistance, Curr. 

Opin. Genet. Dev., 6, 610, 1996. 


462 


19. Narasaki, F., Matsuo, I., Ikuno, N., Fukuda, M., Soda, H., and Oka, M., Multidrug resistance-associated 

protein (MRP) gene expression in human lung cancer, Anticancer Res., 16, 2079, 1996. 

20. Zaman, G. J., Lankelma, J., van Tellingen, Q., Beijnen, J., Dekker, H., Paulusma, C., Oude Elferink, 

R. P., Baas, F., and Borst, P., Role of gluthathione in the export of compounds from cells by the 

multidrug resistance-associated protein, Proc. Natl Acad. Sci. USA, 92, 7690, 1995. 

21. De Jong, S., Zijlstra, J. G., de-Vries, E. G., and Mulder, N. H., Reduced DNA topoisomerase II activity 

and drug-induced DNA cleavage activity in an adriamycin-resistant human small cell lung carcinoma 

cell line, Cancer Res., 50, 304, 1990. 

22. Sugawara, I., Akiyama, S., Scheper, R. J., and Itoyama, S., Lung resistance protein (LRP) expression 

in human normal tissues in comparison with that of MDR1 and MRP, Cancer Lett., 112, 23, 1997. 

23. Pohl, G., Filipits, M., Suchomel, R. W., Stranzl, T., Depisch, D., and Pirker, R., Expression of the lung 

resistance protein (LRP) in primary breast cancer, Anticancer Res., 19, 5051, 1999. 

24. William, S. D. and Scheper, R. J., Lung resistance-related protein: Determining its role in multidrug 

resistance, J. Natl Cancer Inst., 91, 1604, 1999. 

25. Gonçlaves, A., Braguer, D., Kamath, K., Martello, L., Briand, C., Horwitz, S., Wilson, L., and Jordan, 

M. A., Resistance to taxol in lung cancer cells associated with increased microtubule dynamics, Proc. 


26. Rogan, A. M., Hamilton, T. C., Young, R. C., Klecker, R. W., and Ozols, R. F., Reversal of adriamycin 

resistance by verapamil in human ovarian cancer, Science, 224, 994, 1984. 

27. Germann, U. A., Shlyakhter, D., Mason, V. S., Zelle, R. E., Duffy, J. P., Gallulo, U., Armistead, D. M., 

Saunders, J. O., Boger, J., and Harding, M. W., Cellular and biochemical characterization of VX-710 

as a chemosensitizer: Reversal of P-glycoprotein-mediated multidrug resistance in vitro, Anticancer 

Drugs, 8, 125, 1997. 

28. Dale, I. L., Tuffley, W., Callaghan, R., Holmes, J. A., Martin, K., Luscombe, M., Mistry, P., Ryder, 

H., Stewart, A. J., Charlton, P., Twentyman, P. R., and Bevan, P., Reversal of P-glycoproteinmediated 

multidrug resistance by XR9051, a novel diketopiperazine derivative, Br. J. Cancer, 78, 

885, 1998. 

29. Fisher, G. A., Lum, B. L., Hausdorff, J., and Sikic, B.I, Pharmacological considerations in the 

modulation of multidrug resistance, Eur. J. Cancer, 32A, 1082, 1996. 

30. Mistry, P., Stewart, A. J., Dangerfield, W., Okiji, S., Liddle, C., Bootle, D., Plumb, J. A., Templetion, 

D., and Charlton, P., In vitro and in vivo reversal of P-glycoprotein-mediated multidrug resistance by 

a noel potent modulator, XR9576, Cancer Res., 61, 749, 2001. 

31. Simon, S. M., Role of organelle pH in tumor cell biology and drug resistance, Drug Discov. Today,4, 

32, 1999. 

32. Busa, W. B. and Nuccitelli, R., Metabolic regulatoin via intracellular pH, Am. J. Physiol., 246, R409, 

1984. 

33. Simon, S., Roy, D., and Schindler, M., Intracellular pH and the control of multidrug resistance, Proc. 


34. Belhoussine, R., Morjani, H., Millot, J. M., Sharonov, S., and Manfait, M., Confocal scanning 

microspectrofluorometry reveals specific anthracycline accumulation in cytoplasmic organelles of 

multidrug-resistant cancer cells, J. Histochem. Cytochem., 46, 1369, 1998. 

35. Tang, F., Horie, K., and Borchardt, R. T., Are MDCK cells transfected with the human MDR1 gene a 

good model of the human intestinal mucosa?, Pharm. Res., 19, 765, 2002. 

36. Lee, E. S., Shin, H. J., Na, K., and Bae, Y. H., Poly(L-histidine)-PEG block copolymer micelles and 

pH-induced destabilization, J. Control Release, 90, 363, 2003. 

37. Patchornik, A., Berger, A., and Katchalski, E., Poly-L-histidine, J. Am. Chem. Soc., 79, 5227, 1957. 

38. Urry, D. W., Peng, S., Gowda, D. C., Parker, T. M., and Harris, R. D., Comparison of electrostaticand 

hydrophobic-induced pKa shifts in polypentapeptides. The lysine residue, Chem. Phys. Lett., 225, 

97, 1994. 

39. Marques, C. M., Bunchy micelles, Langmuir, 13, 1430, 1997. 

40. La, S. B., Okano, T., and Kataoka, K., Preparation and characterization of the micelle-forming polymeric 

drug indomethacin-incorporated poly(ethylene oxide)-poly(beta-benyzl-L-aspartate) block 

copolymer micelle, J. Pharm. Sci., 85, 85, 1996. 

41. Lee, E. S., Na, K., and Bae, Y. H., Polymeric micelle for tumor pH and folate mediated targeting, 

J. Control Release, 91, 103, 2003. 



42. Ning, Z., Da, D., Rudoll, T. L., Needham, D., Whorton, A. R., and Dewhirst, M. W., Increased 

microvascular permeability contributes to preferential accumulation of stealth liposomes in tumor 

tissue, Cancer Res., 53, 3764, 1993. 


effect in macromolecular therapeutics: A review, J. Control Release, 65, 271, 2000. 



xenograft, Cancer Res., 54, 4564, 1994. 

45. Oku, N., Namba, Y., and Okada, Y., Tumor accumulation of novel RES-avoiding liposomes, Biochim. 

Biophys. Acta., 1126, 225, 1992. 

46. Hobbs, S. K., Monsky, W. L., Yuan, F., Roberts, W. G., Griffith, L., Torchilin, V. P., and Jain, R. K., 

Regulation of transport pathways in tumor vessels: Role of tumor type and microenvironment, Proc. 


47. Charrois, G. J. R. and Allen, T. M., Rate of biodistribution of STEALTH w liposomes to tumor and 

skin: Influence diameter and implications for toxicity and therapeutic activity, Biochim. Biophys. 

Acta., 1609, 102, 2003. 

48. Takeuchi, H., Kojima, H., Toyoda, T., Yamamoto, H., Hinox, T., and Kawashima, Y., Prolonged 

circulation time of doxorubicin-loaded liposomes coated with a modified polyvinyl alcohol after 

intravenous injection in rats, Eur. J. Pharm. Biopharm., 48, 123, 1999. 

49. Goren, D., Horowitz, A. T., Tzemach, D., Tarshish, M., Zalipsky, S., and Gabizon, A., Nuclear 

delivery of doxorubicin via folate-targeted liposomes with bypass of multidrug-resistantce efflux 

pump, Clin. Cancer Res., 6, 1949, 2000. 

50. Shiah, J. G., Dvorak, M., Kopeckova, P., Sun, Y., Peterson, C. M., and Kopecek, J., Biodistribution 

and antitumour efficacy of long-circulating N-(2-hydroxypropyl) methacrylamide copolymer-doxorubicin 

conjugates in nude mice, Eur. J. Cancer, 37, 131, 2001. 

51. Lee, E. S., Na, K., and Bae, Y. H., Doxorubicin loaded pH-sensitive polymeric micelles for reversal of 

resistant MCF-7 tumor, J. Control Release, 103, 405, 2005. 

52. Benns, J. M., Choi, J. S., Mahato, R. I., Park, J. S., and Kim, S. W., pH-sensitive cationic polymer gene 

delivery vehicle: N-Ac-poly(L-histidine)-graft-poly(L-lysine) comb shaped polymer, Bioconjug. 

Chem., 11, 637, 2000. 

53. Na, K. and Bae, Y. H., Self-assembled hydrogel nanoparticles responsive to tumor extracellular pH 

from pullulan derivative/sulfonamide conjugate: Characterization, aggregation and adriamycin 

release in vitro, Pharm. Res., 19, 681, 2002. 

54. Gottesman, M. M. and Pastan, I., Biochemistry of multidrug resistance mediated by the multidrug 

transporter, Ann. Rev. Biochem., 62, 385, 1993. 

55. Thiebaut, F., Tsuruo, T., Hamada, H., Gottesman, M. M., Pastan, I., and Willingham, M. C., Cellular 

localization of the multidrug-resistance gene product P-glycoprotein in normal human tissues, Proc. 


56. Baldini, N., Scotlandi, K., Serra, M., Shikita, T., Zini, N., Ognibene, A., Santi, S., Ferrachini, R., and 

Maraldi, N. M., Nuclear immunolocalization of P-glycoprotein in multidrug-resistant cell lines 

showing similar mechanisms of doxorubicin distribution, Eur. J. Cell Biol., 68, 226, 1995. 

57. Gottesman, M. M., Pastan, I., and Ambudkar, S. V., P-glycoprotein and multidrug resistance, Curr. 

Opin. Genet. Dev., 6, 610, 1996. 

58. Mechetner, E., Kyshtoobayeva, A., Zonis, S., Kim, H., Stroup, R., Garcia, R., Parker, R. J., and 

Fruehauf, J. P., Levels of multidrug resistance (MDR1) P-glycoprotein expression by human 

breast cancer correlate with in vitro resistance to taxol and doxorubicin, Clin. Cancer Res., 4, 

389, 1998. 

59. Mahoney, B. P., Raghunand, N., Baggett, B., and Gillies, R. J., Tumor acidity, ion trapping and 

chemotherapeutics. I. Acid pH affects the distribution of chemotherapeutic agents in vitro, Biochem. 

Pharmacol., 66, 1207, 2003. 

60. Murthy, N., Robichaud, J. R., Tirrell, D. A., Stayton, P. S., and Hoffman, A. S., The design and 

synthesis of polymers for eukaryotic membrane disruption, J. Control Release, 61, 137, 1999. 

61. van Sluis, R., Bhujwalla, Z. M., Raghunand, N., Ballesteros, P., Alvarez, J., Cerdán, S., Galons, J. P., 

and Gillies, R. J., In vivo imaging of extracellular pH using 1 H MRSI, Magn. Reso. Med., 41, 743, 

1999. 


464 


62. Goren, D., Horowitz, A. T., Tzemach, D., Tarshish, M., Zalipsky, S., and Gabizon, A., Nuclear 

delivery of doxorubicin via folate-targeted liposomes with bypass of multidrug-resistantce efflux 

pump, Clin. Cancer Res., 6, 1949, 2000. 

63. Putnam, D., Gentry, C. A., Pack, D. W., and Langer, R., Polymer-based gene delivery with low 

cytotoxicity by a unique balance of side-chain termini, Proc. Natl Acad. Sci., 98, 1200, 2001. 

64. Sabharanjak, S. and Mayor, S., Folate receptor endocytosis and trafficking, Adv. Drug Deliv. Rev., 56, 

1099, 2004. 


23 

CONTENTS 

cRGD-Encoded, MRI-Visible 

Polymeric Micelles for 

Tumor-Targeted Drug Delivery 

Jinming Gao, Norased Nasongkla, and 

Chalermchai Khemtong 

23.1 Introduction ....................................................................................................................... 465 

23.2 Integrin avb3-Mediated Drug Targeting to Angiogenic Tumor Vasculature .................. 466 

23.3 cRGD-Encoded Micelles Facilitate Doxorubicin Delivery to 

Tumor Endothelial Cells ................................................................................................... 467 

23.4 cRGD-Encoded Micelles Increase Doxorubicin Toxicity over 

Non-cRGD Micelles .......................................................................................................... 468 

23.5 Non-Invasive Imaging Methods for Pharmacokinetic Studies......................................... 469 

23.6 SPIO as MR T2 Contrast Agent ....................................................................................... 470 

23.7 Development of MRI-Visible Micelles ............................................................................ 470 

23.8 Non-Invasive Monitoring of Cell and Tumor Targeting by 

cRGD-Encoded, SPIO-Loaded Micelles .......................................................................... 472 

23.9 Conclusions ....................................................................................................................... 473 

Acknowledgments ......................................................................................................................... 474 

References ..................................................................................................................................... 474 


Remarkable progress has been made in recent years to establish polymer micelles as a novel 

multifunctional platform of nanoscale constructs for diagnostic and therapeutic applications. 1–5 

Polymer micelles are composed of amphiphilic block copolymers that contain distinguished hydrophobic 

and hydrophilic segments. The distinct chemical nature of the two blocks results in 

thermodynamic phase separation in aqueous solution and formation of nanoscopic supramolecular 

core/shell structures. During the micellization process, the hydrophobic blocks associate to form the 

core region, whereas the hydrophilic blocks form the shell that separates the core from the aqueous 

medium. This unique architecture enables the micelle core to serve as a nanoscopic depot for 

therapeutic/contrast agents and the shell as biospecific surfaces for targeting applications 

(Figure 23.1). 

Biocompatible and biodegradable block copolymers are currently used in micelle design. For 

the hydrophilic segment, poly(ethylene glycol) (PEG) blocks with a molecular weight of 2–15 kD 

are most often used. Meanwhile, the hydrophobic blocks include a variety of polymers such as 


465

466 

Hydrophobic 

core 

Hydrated 

shell 

∅=20 

–200 nm 

FIGURE 23.1 Micellar nanomedicine platform for targeted drug delivery. 


Targeting ligand 

Iron oxide 

Amphiphilic polymer: 

Hydrophoilic block 

Hydrophobic block 

poly(D, L-lactide), poly(3-caprolactone), poly(b-benzoyl-L-aspartate), and poly(g-benzyl-Lglutamate). 

1,5 Different hydrophobic blocks can be used to provide different hydrophobic environments 

(e.g., aromatic vs. aliphatic) to maximize interactions with encapsulated agents. 

Pioneering work from several leading groups has significantly advanced polymeric micelles as 

versatile carrier systems for the delivery of hydrophobic drugs. For instance, Kataoka’s group has 

recently developed pH-sensitive micelles by attaching hydrophobic doxorubicin (DOX) to the 

polymer backbone via a hydrazone linkage. This acid-labile linkage undergoes cleavage at endosomal/lysosomal 

pH to release the conjugated DOX that leads to significantly improved anti-tumor 

efficacy. 6,7 A powerful means to improve the delivery efficiency of micelles is to introduce 

targeting ligands on the micelle surface to achieve an active targeting function. Among the yet 

limited literature reports dedicated to the active targeting of polymer-related micelle systems, 

recent work by Torchilin’s group provides a landmark study in this direction. 8 Torchilin and 

coworkers have introduced two kinds of antibodies, mAb 2C5 and mAb 2G4, to the distal PEG 

end of phosphatidylethanolamine–poly(ethylene glycol) (PEG–PE) micelles. Not only can the 

immunomicelles with cancer-specific 2C5 antibody specifically bind to cancer cells in vitro, 

they can also increase drug accumulation in Lewis lung carcinoma tumors in mice. When compared 

to administration with either taxol or non-targeting micelles, immunomicelles demonstrate the highest 

efficacy of tumor growth inhibition. These advances bring out the unprecedented promise that polymeric 

micelles can serve as effective carriers for a large variety of poorly soluble drugs. 

23.2 INTEGRIN avb3-MEDIATED DRUG TARGETING TO ANGIOGENIC 

TUMOR VASCULATURE 

Since the original report by Cheresh and coworkers, 9 numerous studies have demonstrated that avb3 

integrin is an important receptor affecting growth, local invasiveness, and metastatic potential of 

tumors. 10–12 In general, integrins are a family of heterodimeric membrane receptors that bind to 

cell-surface adhesion molecules and extracellular matrix proteins (ECM), and they play a critical 

role in tissue morphogenesis, repair, and remodeling. The function of integrins during angiogenesis 

has been most extensively studied with a vb 3 that is not readily detectable in quiescent vessels but 

becomes highly expressed in angiogenic vessels. 9 Recently, the crystal structures of the extracellular 

domains of avb3 and avb3/c(-RGDfV-) complex have been determined by Xiong et al. 13,14 

These structures provide molecular insights and better understanding of avb3 receptor-ligand 

interactions and establish the structural basis for future development of more potent and specific 

avb3-binding ligands. 15 

Vascular targeting via avb3-dependent mechanisms offers significant opportunities for the 

development of cancer-targeted therapeutics and diagnostics. For therapeutic applications, 

Ruoslahti and coworkers have conjugated DOX to a cyclic RGD peptide (RGD-4C) as identified 


MRI-Visible Polymeric Micelles 467 

by a phage library. 16 The DOX-RGD-4C conjugates have demonstrated significantly improved 

anti-tumor efficacy in nude mice bearing human breast tumor xenografts. 17 Although quantitative 

pharmacokinetics of DOX-RGD-4C was not reported, it is expected that the renal clearance of the 

peptide-drug conjugate may be fast as a result of its small molecular weight. Recently, Ghanderhari 

and coworkers have shown that RGD-4C-functionalized hydroxypropylmethacrylamide (HPMA) 

polymer led to significantly increased tumor targeting and localization in severe combined 

immunodeficiency disease mice bearing DU145 or PC-3 prostate tumor xenografts over the 

RGE-4C-HPMA control. 18,19 This work opens up many exciting opportunities to broaden 

HPMA scaffold in targeted drug delivery applications. Recently, Cheresh and coworkers have 

reported the success of a vb 3-targeted gene therapy of cancer with cationic polymerized lipidbased 

nanoparticles. A mutant raf gene was delivered to tumor endothelial cells that led to 

tumor cell apoptosis and sustained regression of primary and metastatic melanomas in mice. 20 

These data provide useful precedence for introducing avb3-targeted ligands on the nano delivery 

systems to enhance their targeting efficiency to tumor vasculature. 

23.3 CRGD-ENCODED MICELLES FACILITATE DOXORUBICIN DELIVERY 

TO TUMOR ENDOTHELIAL CELLS 

Previous work in this lab has demonstrated the feasibility of producing a cyclic pentapeptide c(Arg- 

Gly-Asp-d-Phe-Lys, cRGDFK—also referred to as cRGD herein)-encoded polymeric micelles and 

micelle targeting to avb3-overexpresing tumor endothelial cells (SLK cells). 21 Figure 23.2 

illustrates the schematic diagram of synthesis of maleimide–terminated poly(ethylene glycol)– 

poly(3-caprolactone) (MAL–PEG–PCL) copolymer and cRGD attachment on the micelle 

surface. Doxorubicin was also encapsulated inside the micelle core. The intrinsic fluorescent 

properties of doxorubicin (lexZ485 nm, lemZ595 nm) permit the study of cell uptake by flow 

cytometry and confocal laser scanning microscopy. 

Figure 23.3a shows that the percentage of cell uptake increased with increasing cRGD density 

on the micelle surface. With 5% cRGD surface density, a modest 3-fold increase of cell uptake was 

O 

O 

O HO-PEG-N 

O 

ε-caprolactone HO-PEG-MAL 

Δ 

Sn(Oct) 2 

O H 

H H 

N 

HO 

HN 

O 

HN 

H 

O 

O 

N 

O 

cRGD-SH 

O 

O 

H-[OCH2CH2CH2CH2CH2C] m-(O-CH2CH2 ) n-N PCL-b-PEG-MAL 

SPIO 

Water and sonication 

cRGD functionalized SPIO-loaded micelle Maleimide containing micelle 

O 

NH 

SH 

O 

NH2 NH 

NH 

FIGURE 23.2 Synthesis of MAL–PEG–PCL copolymer and preparation of cRGD-encoded, DOX-loaded 

micelles. (From Nasongkla, N., Shuai, X., Ai, H., Weinberg, B. D., Pink, J., Boothman, D. A., and Gao, J., 

Angew. Chem. Int. Ed. Engl., 43, 6323, 2004. With permission.) 


O

468 

Percentage of cell uptake 

(a) 

100 

80 

60 

40 

20 

0 

0 5 16 76 76+ 

Micelle composition (% cRGD) free RGD 

(b) 20 μm (c) 

20 μm 

FIGURE 23.3 (See color insert following page 522.) (a) Percentage of micelle uptake in SLK tumor 

endothelial cells by flow cytometry as a function of cRGD density (0–76%) on the micelle surface. The last 

panel shows that the cell uptake of 76% cRGD-micelles is inhibited by the presence of free RGD ligands 

(9 mM) in solution. (b, c) Confocal laser scanning microscopy images of SLK cells treated with 0% (b) and 

16% (c) cRGD-micelles after incubation for 2 h. The scale bars are 20 mm in both images. (From Nasongkla, 

N., Shuai, X., Ai, H., Weinberg, B. D., Pink, J., Boothman, D. A., and Gao, J., Angew. Chem. Int. Ed. Engl., 43, 

6323, 2004. With permission.) 

observed. A 30-fold increase was observed by flow cytometry with 76% cRGD attachment. In the 

presence of excess free RGD ligands, the avb3-mediated cell uptake can be completely inhibited 

(last panel, Figure 23.3a). Confocal laser scanning microscopy further supports the cRGDenhanced 

cell uptake (Figure 23.3b and c) and demonstrates that a majority of the micelles were 

taken up via a vb 3-mediated endocytosis and localized mainly in the cytoplasmic compartments in 

the cell (as shown by the red punctated dots). Cell nuclei were stained blue by Hoechst 33342 

(lexZ352 nm, lemZ455 nm) and overlaid with doxorubicin fluorescent images (lexZ485 nm, 

lemZ595 nm). 

23.4 CRGD-ENCODED MICELLES INCREASE DOXORUBICIN 

TOXICITY OVER NON-CRGD MICELLES 


Figure 23.4 shows the growth inhibition of SLK cells by different micelle formulations. DOX 

concentration was kept the same at 1 mM for all samples. In these experiments, different DOX 



Percentage of growth inhibition 

140 

120 

100 

80 

60 

40 

20 

0 

Free DOX 0% RGD 5% RGD 16% RGD 16% RGD+ 

16% RGD 

RGD w/o DOX 

Micelle formulation 

FIGURE 23.4 Cell growth in SLK cells for different compositions of SPIO-loaded micelles at 1 mM DOX 

concentration. The percentage of growth was calculated as the ratio of cell numbers for treated samples over 

untreated control. Error bars were obtained from triplicate samples. 

samples were exposed to SLK cells for 1 h. After 1 h, the DOX-containing media were replaced 

with cell culture media and kept for an additional four days until the cells in the untreated control 

group reached confluence. Percentage of growth (PG) was calculated as the ratio of the number of 

cancer cells of the treated group over the untreated control. At 1 mM DOX dose, free drug demonstrated 

the lowest growth at 24.4G2%. For micelle-delivered DOX, growth decreased with an 

increase of cRGD density on the micelle surface. The PG values for 0%, 5%, and 16% cRGDmicelles 

were 92G8%, 85G10%, and 51G2.58%, respectively. The cRGD micelle toxicity was 

inhibited by the free RGD ligand co-incubated with the micelles (PG 78G10%, Figure 23.4). In 

addition, cRGD-encoded, DOX-free micelles did not show much tumor growth inhibition (115G 

11%) at the same micelle dose. These data demonstrate the success of cRGD-encoded micelles in 

enhancing the drug targeting efficiency and cytotoxic response in avb3-overexpresing SLK cells. 

23.5 NON-INVASIVE IMAGING METHODS FOR PHARMACOKINETIC 

STUDIES 

Conventional clinical pharmacokinetic studies require the collection of blood or tissue samples to 

assess the clearance rate, tissue distribution, and concentration vs. time relationships of 

chemotherapy. Such procedures are invasive and not compliant with patients especially when 

multiple samples need to be collected at different time points. In addition, the quantitative interpretation 

of these data is complicated by individual variability and uncertainty in the technical 

procedure (e.g., extraction of drugs from tissue may not be complete). 

To circumvent these limitations, significant research efforts have focused on the development 

of non-invasive imaging techniques to characterize the drug pharmacokinetics in the target tissue 

in vivo. 22–25 Such imaging techniques include single photon emission computed tomography 

(SPECT), 26,27 positron-emission tomography (PET, dual photons), 26,27 and magnetic resonance 

imaging (MRI). 28 SPECT imaging requires the use of radionuclides (e.g., 111 In, 99m Tc) that emit 

high-energy gamma photons (typically 60–600 keV) whereas PET detects positron annihilation 

radiation (511 keV). Although these two methods provide the most sensitive measurement of 

radiolabeled drugs (10 K10 –10 K12 mol), they have very low spatial resolutions. For example, the 

resolutions of small animal SPECT and PET instruments are approximately 2 and 5 mm, respectively. 

26,27 Such resolution is not adequate to accurately assess drug distribution within tumor tissue, 

especially in a small animal tumor model (e.g., tumor xenograft in mice). 


470 

MRI is another imaging modality that can be used to study in situ pharmacokinetics. 28 MRI 

detects the magnetization changes of nuclei that have net angular momentum such as 1 H, 13 C, and 

19 F. Examples of drugs studied by MRI include 5-fluorouracil ( 19 F), 13 C-labelled temozolomide, 

and ifosfamide ( 31 P). MRI has much higher resolution (w100 mm) than nuclear imaging methods. 

Its main limitation is low sensitivity, making it difficult to detect drugs at their therapeutic concentrations. 

28 Use of exogenous MR contrast agents (e.g., SPIO) can dramatically increase the 

sensitivity of MR detection. 

23.6 SPIO AS MR T2 CONTRAST AGENT 

Superparamagnetic iron oxide (SPIO) nanoparticles have been extensively studied in the past 

decade for their use as MR contrast agents. Unlike the low molecular weight, paramagnetic 

metal chelates such as Gd-DTPA (T1 contrast agent) and SPIO nanoparticles are considered T2negative 

contrast agents with substantially higher T2 and T1 relaxivity compared to T1 agents. 29 

SPIO agents are composed of iron oxide nanocrystals with the general formula Fe 3C 

2 O 3M 2C O, 

where M 2C is a divalent metal ion such as iron, manganese, nickel, cobalt, or magnesium. 

When M 2C is ferrous iron (Fe 2C ), SPIO becomes magnetite. In the absence of an external magnetic 

field, the magnetic domains inside SPIO are randomly oriented with no net magnetic field. An 

external magnetic field can cause the magnetic dipoles of the magnetic domains to reorient, leading 

to dramatically increased magnetic moments and significantly shortened relaxations in both T1 and 

T2/T2* relaxation processes. 29 

SPIO nanoparticles have been classified into four different categories based on their sizes and 

applications: large, standard, ultrasmall, and monocrystalline agents. Large SPIO agents (O200 nm) 

are mainly used for oral gastrointestinal lumen imaging. For example, AMI-121 (Advanced 

Magnetics) is a clinically approved T2 agent with a hydrodynamic diameter of 300 nm and 

T2andT1relaxivitiesof72and3.2FemM K1 s K1 , respectively. 30 Standard SPIO agents 

(50–200 nm) are easily sequestered by the RES in the liver and spleen upon intravenous injection 

and are primarily used for liver and spleen imaging. The iron oxide crystal (4.8–5.6 nm) in AMI-25 

(Endorem w by Guerbet and Feridex w by Berlex Lab), a clinically approved agent, is coated with 

dextran to yield a final hydrodynamic diameter between 80 and 150 nm. The T2 and T1 relaxivities 

are 98.3 and 23.9 Fe mM K1 s K1 , respectively. 31 AMI-25 efficiently accumulates in the liver (w80% 

of injected dose) and spleen (5–10%) within minutes of administration, and it has a short blood 

half-life (6 min). 

Weissleder and coworkers have pioneered the development of ultrasmall (USPIO) and monocrystalline 

iron oxide nanoparticles (MION). 32–35 Two USPIO agents, AMI-227 and CLARISCAN, 

are currently in phase III clinical trials. AMI-227 is composed of a crystalline core of 4–6 nm 

surrounded by a dextran coating. Its hydrodynamic diameter is 20–40 nm, and it has T2 and T1 

relaxivities of 44.1 and 21.6 Fe mM K1 s K1 , respectively. 36 Because of its smaller size, AMI-227 

showed an increased blood half-life over 24 h. 37 Consequently, AMI-227 can be used as a blood 

pool agent for MR angiography. Eventually, these particles are taken up by the RES of lymph nodes 

and become particularly useful for MR lymphography. 38,39 

23.7 DEVELOPMENT OF MRI-VISIBLE MICELLES 


Recently, a new class of MRI-visible nanoparticles was developed through the incorporation of 

hydrophobic SPIO nanoparticles inside polymer micelles. 40 High loading of SPIO was achieved 

inside the micelle core that led to a significant increase in T2 relaxivity. The resulting SPIO-loaded 

micelles have shown superb sensitivity under a clinical 1.5 T MR scanner. 

First, monodisperse SPIO (Fe 3O 4) nanoparticles with different sizes were synthesized 

following published work by Sun and coworkers. 41 Transmission electron microscopy (TEM) 



(a) 

(b) 

(c) 

Volume 

Volume 

120 

100 

80 

60 

40 

20 

(d) 

Volume 

0 

11 

120 

100 

80 

60 

40 

20 

0 

11 

(e) 

(f) 

120 

100 

80 

60 

40 

20 

0 

11 

17 26 41 62 89 119 

Diameter (nm) 

17 26 41 60 80 108 

Diameter (nm) 

17 26 41 60 80 110 

Diameter (nm) 

FIGURE 23.5 TEM images and DLS data of different SPIO-micelle formulations. (a, d) Single SPIO (4 nm in 

diameter) encapsulated by PEG5k–DSPE lipid; (b, e) SPIO (4 nm)-loaded PEG5k-b-PCL5k micelles; (c, f) 

SPIO (8 nm)-loaded PCL5k-b-PEG5k micelles. A cluster of SPIO particles are loaded inside each PCL5k-b- 

PEG5k micelle as shown in (b) and (c). (From Ai, H., Flask, C., Weinberg, B., Shuai, X., Pagel, M., Farrell, D., 

Duerk, J., and Gao, J., Adv. Mat., 17, 1949, 2005. With permission.) 

characterization of SPIO nanoparticles showed 4, 8, and 16 nm in diameter. The size distribution is 

uniform for all particle sizes. Distearoylphosphatidyl–ethanolamine–poly(ethylene glycol) 

(PEG5k–DSPE) lipid and amphiphilic diblock copolymer of PCL5k-b-PEG5k were used for the 

micelle formation. Interestingly, PEG5k–DSPE lipid micelles provide individual encapsulation of 

single 4 nm SPIO nanoparticles (Figure 23.5a). In comparison, PEG5k-b-PCL5k polymer micelles 

allow the incorporation of a cluster of 4 and 8 nm SPIO nanoparticles, respectively (Figure 23.5b 

and Figure 23.5c). Dynamic light scattering (DLS) was used to study the size distribution of SPIOloaded 

PEG5k-b-PCL5k polymer micelles. DLS experiments were performed using the micelle 

solutions with 908 scattering angle at 258C. Figure 23.5d shows a representative DLS diagram of 

PEG5k–DSPE micelles containing 4 nm SPIO nanoparticles. The mean hydrodynamic diameter 

was measured to be 17G1 nm. The mean hydrodynamic diameters were 75G4 and 97G6 nm for 


472 

4 nm SPIO 

DSPE-PEG5k 

4 nm SPIO 

PCL5k-PEG5k 

Fe Concentration (μM) 

1040 520 260 130 65 32.5 16.25 H 2O 


FIGURE 23.6 T2-weighted MRI scan (TRZ5000 ms, TEZ40 ms) of single 4 nm SPIO encapsulated by 

PEG5k-DSPE and a cluster of 4 nm SPIO loaded in PEG5k-PCL5k micelles. MRI intensity at the same Fe 

concentration is compared. (From Ai, H., Flask, C., Weinberg, B., Shuai, X., Pagel, M., Farrell, D., Duerk, J., 

and Gao, J., Adv. Mat., 17, 1949, 2005. With permission.) 

PEG-b-PCL micelles containing 4 and 8 nm SPIO particles (Figure 23.5e and Figure 23.5f), 

respectively. 

The T1 and T2 relaxivities of different micelle particles were determined on a clinical 1.5 T 

scanner (Siemens Sonata). The T1 relaxivities for different micelle formulations were comparable 

and ranged from 1.3 to 2.9 Fe mM K1 s K1 . In contrast, the T2 relaxivities of micelles notably 

depended on SPIO clustering, SPIO diameter, and loading density. Figure 23.6 illustrates the 

MR signal intensity as a function of Fe concentration for two micelle formulations. The top row 

shows SPIO (4 nm) particles individually encapsulated by PEG5k–DSPE lipid (sample in 

Figure 23.5a), and the bottom row shows PEG-b-PCL micelles loaded with the same 4 nm SPIO 

particles (sample in Figure 23.5b). At the same Fe concentration, SPIO-loaded PEG-b-PCL 

micelles showed significantly increased contrast in these T2-weighted images. The T2 relaxivity 

increased from 25 for single SPIO micelles to 169 Fe mM K1 s K1 for SPIO-loaded polymeric 

micelles (w7 fold increase). Clustering of SPIO particles considerably increased the T2 relaxivity 

per Fe concentration. For polymeric micelles, further increase in SPIO diameter (e.g., from 4 to 8 

and 16 nm) also led to significantly increased loading density and enhanced T2 relaxivity. The 

values of T2 relaxivity reached 318 and 471 Fe mM K1 s K1 for micelles loaded with 8 and 16 nm 

Fe3O4 particles, respectively. 40 

The MR sensitivity of detection was measured for the micelle formulations using a T2weighted 

scan (TRZ5000 ms, TEZ40 ms) in an in vitro test tube experiment. The sensitivity 

limit is defined as the micelle concentration where the MR signal intensity decreases to 50% of the 

signal of pure water. For PEG-b-PCL micelles loaded with 16 nm SPIO particles, the highest 

sensitivity limit was observed at 5.2 mg/mL. In this micelle formulation, the micelle diameter is 

110G9 nm, and Fe3O4 loading is 54.2%. Because the molecular weight of typical micelles is on the 

order of 10 6 g/mol, 2 the above sensitivity limit corresponds to a micelle concentration of approximately 

5 nM. 

23.8 NON-INVASIVE MONITORING OF CELL AND TUMOR TARGETING 

BY CRGD-ENCODED, SPIO-LOADED MICELLES 

Recently, cRGD-encoded, SPIO-loaded polymer micelles using similar procedures were successfully 

developed as shown in Figure 23.2. In this system, 8 nm SPIO nanoparticles were 

encapsulated (6.7 wt.%) into cRGD-encoded (50%) and non-cRGD (0%) PEG–PLA micelles. 

The mean hydrodynamic diameters of both micelle formulations are approximately 40 nm by 

dynamic light scattering measurement. In cell uptake experiments, tumor SLK endothelial cells 

(1.2!10 6 cells) were co-incubated with SPIO micelles at the final iron concentrations of 0, and 

12.5 mg/mL in 10% fetal bovine serum cell culture medium for two hours. The medium was then 

removed, and the cells were gently washed twice with HBSS. Cells were detached by trypsin, and 



MRI Intensity 

(a) 

15 

10 

cell pellets were collected after centrifugation at 1200 rpm for 5 min. Cells were then fixed 

with 2% paraformaldehyde in phosphate-buffered saline (PBS) (200 mL) for two hours and 

concentrated by centrifugation again. Finally, each sample was dispersed in 20 mL of 2% 

agarose gel in PBS. A 384-well plate was used to contain the samples for MRI imaging. MR 

images were collected with conventional T2-weighted spin echo acquisition parameters (TRZ 

6,000 ms, TEZ90 ms). 

Figure 23.7 shows the MR signal intensity of SLK cells incubated with cRGD-encoded and 

non-cRGD micelles. Three major conclusions can be drawn from these experiments. First, cRGDencoded 

micelles led to significant reduction of MR signal amplitude over non-cRGD micelles as a 

result of the increased micelle uptake in SLK cells. Second, non-specific uptake of non-cRGD 

micelles is relatively small with minimal change in MR intensity over the Fe concentration range 

from 0 to 12.5 mg/mL. Third, MR detection is highly sensitive in imaging tumor cells (1.2!10 6 

SLK cells at 12!10 6 cells/mL) with specific avb3-mediated uptake of Fe3O4 micelles. At 12.5 

Fe mg/mL, MR amplitude decreased from 9.44 for non-cRGD micelles to 4.98 for cRGD-encoded 

micelles. The above data demonstrate the feasibility in producing biologically specific, highly 

sensitive MR molecular probes based on the micelle construct and design. 


5 

0 

50 % 

0 % 

12.5 mg/mL SPIO Cell PBS 

FIGURE 23.7 (a) MRI signal intensity of SLK cells treated with SPIO-loaded PEG–PLA micelles. SLK cells 

with micelle treatment and PBS buffer were used as control. Images (b) and (c) correspond to the cells 

incubated with 50% and 0% cRGD-micelles at 12.5 mg Fe/mL, respectively. (d) MR image of the vial 

containing SLK cells without any micelle incubation. (e) MR image of the vial containing PBS. MRI intensity 

is significantly darkened in cells incubated with cRGD-encoded micelles (b). 

The long-term goal of this research is to establish polymer micelles as a safe and efficacious 

nanomedicine platform for targeted cancer therapy. The cRGD-encoded micelles have shown 

superb targeting efficiency in cell culture in vitro that establishes the micelle foundation for 

future in vivo evaluations in tumor-bearing mice. To facilitate the evaluation of tumor targeting 

efficiency of these micelles, MRI-visible micelles have been developed. Clustering of SPIO 

particles inside the micelle core led to a dramatic increase in T2 relaxivity. Together with high 

loading density of Fe3O4 inside polymer micelles, they led to an ultra sensitivity of detection by 

MRI per micelle basis (w5 nM). The much improved sensitivity of SPIO-loaded micelles may 

prove essential in detecting low concentrations of micelle nanoparticles for in vivo intratumoral 

tissue distribution studies by MRI. 


(b) 

(c) 

(d) 

(e)

474 


This work was supported by the National Institutes of Health grants R01 CA90696 and R21 

EB005394 to Jinming Gao. Norased Nasongkla acknowledges the Royal Thai Government for a 

predoctoral fellowship support. 

REFERENCES 


1. Adams, M. L., Lavasanifar, A., and Kwon, G. S., Amphiphilic block copolymers for drug delivery, 

J. Pharm. Sci., 92, 1343–1355, 2003. 

2. Kwon, G. S., Polymeric micelles for delivery of poorly water-soluble compounds, Crit. Rev. Ther. 

Drug. Carrier Syst., 20, 357–403, 2003. 

3. Lukyanov, A. N. and Torchilin, V. P., Micelles from lipid derivatives of water-soluble polymers as 

delivery systems for poorly soluble drugs, Adv. Drug Deliv. Rev., 56, 1273–1289, 2004. 

4. Otsuka, H., Nagasaki, Y., and Kataoka, K., PEGylated nanoparticles for biological and pharmaceutical 

applications, Adv. Drug Deliv. Rev., 55, 403–419, 2003. 

5. Torchilin, V. P., Targeted polymeric micelles for delivery of poorly soluble drugs, Cell. Mol. Life Sci., 

61, 2549–2559, 2004. 

6. Bae, Y., Fukushima, S., Harada, A., and Kataoka, K., Design of environment-sensitive supramolecular 

assemblies for intracellular drug delivery: Polymeric micelles that are responsive to intracellular 

pH change, Angew. Chem. Int. Ed. Engl., 42, 4640–4643, 2003. 




antitumor efficacy, Bioconjug. Chem., 16, 122–130, 2005. 

8. Torchilin, V. P., Lukyanov, A. N., Gao, Z., and Papahadjopoulos-Sternberg, B., Immunomicelles: 

Targeted pharmaceutical carriers for poorly soluble drugs, Proc. Natl. Acad. Sci. USA, 100, 

6039–6044, 2003. 

9. Brooks, P. C., Clark, R. A., and Cheresh, D. A., Requirement of vascular integrin alpha v beta 3 for 

angiogenesis, Science, 264, 569–571, 1994. 

10. Ruegg, C., Dormond, O., and Foletti, A., Suppression of tumor angiogenesis through the inhibition of 

integrin function and signaling in endothelial cells: Which side to target? Endothelium, 9, 151–160, 2002. 

11. Teti, A., Migliaccio, S., and Baron, R., The role of the alphaVbeta3 integrin in the development of 

osteolytic bone metastases: A pharmacological target for alternative therapy? Calcif. Tissue Int., 71, 

293–299, 2002. 

12. Varner, J. A. and Cheresh, D. A., Tumor angiogenesis and the role of vascular cell integrin alphavbeta3, 

Important Adv. Oncol., 69–87, 1996. 

13. Xiong, J. P., Stehle, T., Diefenbach, B., Zhang, R., Dunker, R., Scott, D. L., Joachimiak, A., 

Goodman, S. L., and Arnaout, M. A., Crystal structure of the extracellular segment of integrin 

alpha Vbeta3, Science, 294, 339–345, 2001. 

14. Xiong, J. P., Stehle, T., Zhang, R., Joachimiak, A., Frech, M., Goodman, S. L., and Arnaout, M. A., 

Crystal structure of the extracellular segment of integrin alpha V beta3 in complex with an Arg–Gly– 

Asp ligand, Science, 296, 151–155, 2002. 

15. Gottschalk, K. E. and Kessler, H., The structures of integrins and integrin-ligand complexes: Implications 

for drug design and signal transduction, Angew. Chem. Int. Ed. Engl., 41, 3767–3774, 2002. 


circulating ligands, Nat. Biotechnol., 15, 542–546, 1997. 


vasculature in a mouse model, Science, 279, 377–380, 1998. 

18. Line, B. R., Mitra, A., Nan, A., and Ghandehari, H., Targeting tumor angiogenesis: Comparison of 

Peptide and polymer-Peptide conjugates, J. Nucl. Med., 46, 1552–1560, 2005. 

19. Mitra, A., Mulholland, J., Nan, A., McNeill, E., Ghandehari, H., and Line, B. R., Targeting tumor 

angiogenic vasculature using polymer-RGD conjugates, J. Control. Release, 102, 191–201, 2005. 

20. Hood, J. D., Bednarski, M., Frausto, R., Guccione, S., Reisfeld, R. A., Xiang, R., and Cheresh, D. A., 

Tumor regression by targeted gene delivery to the neovasculature, Science, 296, 2404–2407, 2002. 



21. Nasongkla, N., Shuai, X., Ai, H., Weinberg, B. D., Pink, J., Boothman, D. A., and Gao, J., cRGDfunctionalized 

polymer micelles for targeted doxorubicin delivery, Angew. Chem. Int. Ed. Engl., 43, 

6323–6327, 2004. 

22. Fischman, A. J., Alpert, N. M., and Rubin, R. H., Pharmacokinetic imaging: A noninvasive method for 

determining drug distribution and action, Clin. Pharmacokinet., 41, 581–602, 2003. 

23. Port, R. E. and Wolf, W., Noninvasive methods to study drug distribution, Invest. New Drugs, 21, 

157–168, 2003. 

24. Seddon, B. M. and Workman, P., The role of functional and molecular imaging in cancer drug 

discovery and development, Br. J. Radiol., 76, S128–S138, 2003. 

25. Singh, M. and Waluch, V., Physics and instrumentation for imaging in-vivo drug distribution, Adv. 

Drug Deliv. Rev., 41, 7–20, 2000. 

26. Bhatnagar, A., Hustinx, R., and Alavi, A., Nuclear imaging methods for non-invasive drug monitoring, 

Adv. Drug Deliv. Rev., 41, 41–54, 2000. 

27. Saleem, A., Aboagye, E. O., and Price, P. M., In vivo monitoring of drugs using radiotracer techniques, 


28. Griffiths, J. R. and Glickson, J. D., Monitoring pharmacokinetics of anticancer drugs: Non-invasive 

investigation using magnetic resonance spectroscopy, Adv. Drug Deliv. Rev., 41, 75–89, 2000. 

29. Wang, Y., Hussain, S., and Krestin, G., Superparamagnetic iron oxide contrast agents: Physicochemical 

characteristics and applications in MR imaging, Eur. Radiol., 11, 2319–2331, 2001. 

30. Hahn, P. F., Stark, D. D., Lewis, J. M., Saini, S., Elizondo, G., Weissleder, R., Fretz, C. J., and 

Ferrucci, J. T., First clinical trial of a new superparamagnetic iron oxide for use as an oral gastrointestinal 

contrast agent in MR imaging, Radiology, 175, 695–700, 1990. 

31. Weissleder, R., Stark, D. D., Engelstad, B. L., Bacon, B. R., Compton, C. C., White, D. L., Jacobs, P., 

and Lewis, J., Superparamagnetic iron oxide: Pharmacokinetics and toxicity, AJR Am. J. Roentgenol., 

152, 167–173, 1989. 

32. Moore, A., Weissleder, R., and Bogdanov, A. Jr., Uptake of dextran-coated monocrystalline iron 

oxides in tumor cells and macrophages, J. Magn. Reson. Imaging, 7, 1140–1145, 1997. 

33. Schaffer, B. K., Linker, C., Papisov, M., Tsai, E., Nossiff, N., Shibata, T., Bogdanov, A. Jr., Brady, 

T. J., and Weissleder, R., MION-ASF: Biokinetics of an MR receptor agent, Magn. Reson. Imaging, 

11, 411–417, 1993. 

34. Weissleder, R., Elizondo, G., Wittenberg, J., Rabito, C. A., Bengele, H. H., and Josephson, L., 

Ultrasmall superparamagnetic iron oxide: Characterization of a new class of contrast agents for 

MR imaging, Radiology, 175, 489–493, 1990. 

35. Weissleder, R., Lee, A. S., Fischman, A. J., Reimer, P., Shen, T., Wilkinson, R., Callahan, R. J., and 

Brady, T. J., Polyclonal human immunoglobulin G labeled with polymeric iron oxide: Antibody MR 

imaging, Radiology, 181, 245–249, 1991. 

36. Jung, C. W. and Jacobs, P., Physical and chemical properties of superparamagnetic iron oxide MR 

contrast agents: Ferumoxides, ferumoxtran, ferumoxsil, Magn. Reson. Imaging, 13, 661–674, 1995. 

37. McLachlan, S. J., Morris, M. R., Lucas, M. A., Fisco, R. A., Eakins, M. N., Fowler, D. R., Scheetz, 

R. B., and Olukotun, A. Y., Phase I clinical evaluation of a new iron oxide MR contrast agent, J. Magn. 

Reson. Imaging, 4, 301–307, 1994. 

38. Anzai, Y., Blackwell, K. E., Hirschowitz, S. L., Rogers, J. W., Sato, Y., Yuh, W. T., Runge, V. M., 

Morris, M. R., McLachlan, S. J., and Lufkin, R. B., Initial clinical experience with dextran-coated 

superparamagnetic iron oxide for detection of lymph node metastases in patients with head and neck 

cancer, Radiology, 192, 709–715, 1994. 

39. Anzai, Y., Brunberg, J. A., and Lufkin, R. B., Imaging of nodal metastases in the head and neck, 

J. Magn. Reson. Imaging, 7, 774–783, 1997. 

40. Ai, H., Flask, C., Weinberg, B., Shuai, X., Pagel, M., Farrell, D., Duerk, J., and Gao, J., Magnetiteloaded 

polymeric micelles as novel magnetic resonance probes, Adv. Mat., 17, 1949–1952, 2005. 

41. Sun, S. and Zeng, H., Size-controlled synthesis of magnetite nanoparticles, J. Am. Chem. Soc., 124, 

8204–8205, 2002. 


24 

CONTENTS 

Targeted Antisense 

Oligonucleotide Micellar 

Delivery Systems 

Ji Hoon Jeong, Sun Hwa Kim, and Tae Gwan Park 

24.1 Introduction ....................................................................................................................... 477 

24.2 Micellar Antisense Oligonucleotide Delivery Systems .................................................... 478 

24.2.1 Poly(D, L-Lactic-Co-Glycolic Acid)-Oligonucleotides 

Hybrid Polymeric Micelles ................................................................................. 478 

24.2.2 Oligonucleotide Delivery Systems Based on Polyelectrolyte 

Complex Micelles ............................................................................................... 478 

24.2.3 In Vivo Applications ........................................................................................... 481 

24.3 Ligand-Mediated Targeting of Antisense Oligonucleotides ............................................ 483 

24.3.1 Ligand-Mediated ODN Targeting Systems Based on PEC Micelles ................ 483 

24.3.2 In Vivo Applications ........................................................................................... 483 

24.4 Conclusions ....................................................................................................................... 484 

References..................................................................................................................................... 484 


Antisense oligonucleotides (antisense ODNs) are receiving increasing attention because of their 

high selectivity toward the sequence of a target mRNA. 1–4 The sequence-specific binding results in 

the degradation of the target mRNA, leading to blockage of the target protein expression. Owing to 

the features, antisense ODNs have been popularly employed for the modulation of specific gene 

expression. The clinical use of antisense ODN, however, is still limited as a result of its intrinsic 

properties: vulnerability to enzymatic digestion by humoral and cellular nucleases, and poor 

cellular uptake in a target tissue. 5–7 The susceptibility of ODN to enzymatic digestion has been 

partly improved by modifying the ODN backbone with either phosphorothioate linkage or amide 

linkage peptide nucleic acid (PNA). 8–10 The problem with poor cellular uptake could be addressed 

by employing cationic lipids and polymers. 11–15 The electrostatic interaction between a cationic 

lipid/polymer and negatively charged DNA or ODN can lead to the formation of polyelectrolyte 

complex (PEC) nanoaggregates that can greatly facilitate cellular uptake of the polyanions by an 

adsorption-induced internalization process called endocytosis. Despite their satisfactory performance 

in intracellular gene transfer in vitro, the use of PEC is still limited because of its 

rapid clearance and undesirable pharmacokinetic profiles upon intravenous administration. 16,17 


477

478 

A series of recent developments for micellar gene delivery systems could provide a chance to 

overcome the potential problems concerning in vivo application of PEC systems. Self-assembling 

polymeric micelles have been considered as potential drug carriers for anti-cancer drugs, including 

nucleic acid-based therapeutic drugs such as antisense ODNs. Most polymeric micelles generally 

have a core-shell structure composed of hydrophobic parts as an internal core and hydrophilic parts 

as a surrounding corona. The hydrophilic corona could protect micelles from nonspecific adsorption 

of serum proteins in the blood stream, leading to an avoidance of opsonization-induced clearance of 

the micelles. The size of polymeric micelles, usually less than 100 nm, provides them with a chance 

to escape from renal exclusion and reticulo-endothelial system (RES). Because of the small size of 

the micelles, the enhanced permeability through loosened endothelial cell junctions in the vicinity 

of solid tumors can also be expected. A hybrid conjugate of ODN with hydrophobic biodegradable 

poly(D, L-lactic-co-glycolic acid) (PLGA) that exhibited micellar properties in aqueous solution 

was recently introduced. 6 PEC micelles are also a new class of gene delivery systems based on 

electrostatically interacted DNA/polymer complexes. The PEC micelles could be prepared from the 

interaction between ODN and cationic block copolymers, including poly(ethylene glycol) (PEG)-bpoly(L-lysine) 

di-block copolymer and PEG-grafted polyethylenimine (PEI). 18,19 Alternatively, the 

PEC micelles could be also formed by interacting PEG-conjugated ODN (ODN–PEG) and cationic 

polymers or peptides. 20 These systems could provide ODN with the enhanced solubility of nanocarriers 

and an improved stability against nuclease attacks. Introduction of cell binding ligands such 

as transferring and folate could also enhance the cellular uptake and improve the in vivo tissue 

distribution of the PEC micelles. 

24.2 MICELLAR ANTISENSE OLIGONUCLEOTIDE DELIVERY SYSTEMS 

24.2.1 POLY(D, L-LACTIC-CO-GLYCOLIC ACID)-OLIGONUCLEOTIDES 

HYBRID POLYMERIC MICELLES 


Micelle-forming amphiphilic copolymers have been utilized as a solubilizer and carrier of various 

hydrophobic drugs, including several water insoluble anti-cancer drugs. An A–B type block copolymer 

of PEG-b-PLGA, containing doxorubicin conjugated to the terminal end of PLGA, could 

form micelles and remarkably improve the cellular uptake of doxorubicin. 21 Slow degradation of 

PLGA chains permit sustained release of the conjugated drug over a desired period. An alternative 

micelle-forming copolymer composed of PLGA and ODN (ODN/PLGA) was synthesized for 

intracellular ODN delivery. 6 Water-insoluble PLGA segments serve as a hydrophobic core and 

hydrophilic ODN as a surrounding corona, resulting in an amphiphilic structure similar to A–B type 

di-block copolymers. The ODN/PLGA could form spherical micelles with a size of ca. 80 nm. The 

micelles have a critical micelle concentration (cmc) ranging from 5 to 10 mg/L, depending on the 

analytical methods. When ODN/PLGA micelles were subjected to incubation in an aqueous 

solution, the micelles released ODN in a sustained manner by controlled degradation of conjugated 

PLGA chains. The micelles demonstrated much higher intracellular delivery efficiency in NIH3T3 

mouse fibroblast cells when compared to naked ODN. The fluorescently labeled micelles were 

distributed over the cytoplasm of the cells, suggesting the micelles were taken up by the cells 

via endocytosis. 

24.2.2 OLIGONUCLEOTIDE DELIVERY SYSTEMS BASED ON POLYELECTROLYTE COMPLEX MICELLES 

Electrostatic interaction between oppositely charged polyelectrolytes (e.g., polycations and DNA) 

leads to the formation of nanoparticular polyelectrolyte complex (PEC) by charge neutralization. 

The nanoparticulates are usually unstable in aqueous milieu as a result of the high incidence of 

collision among particles and their high surface energy. Because they tend to aggregate one 

another, often resulting in precipitation, the complex formation between polycations and 


Targeted Antisense Oligonucleotide Micellar Delivery Systems 479 

polyanions (ODN or plasmid DNA) should be nonstoichiometric; there the surface charge of the 

complexes should be negative (polyanions excess) or positive (polycation excess). The nonstoichiometric 

condition between polycation and polyanion can prevent the premature inter-particular 

aggregation by charge repulsion among the complexes during their formation. 22 However, the 

complexes with negatively charged surfaces are not appropriate for intracellular gene transfer 

because of the negatively charged characteristics of the cellular membrane. The positively 

charged particles are also undesirable for in vivo application because of their susceptibility to 

the nonspecific interaction of serum components such as negatively charged albumin that could 

often lead to inter-particular aggregation in the blood stream or opsonization-mediated clearance 

by reticuloendothelial cells. The introduction of hydrophilic PEG segments to the complexes’ 

surfaces greatly increases their stability, not only enhancing the solubility of the complexes, 

but also efficiently preventing inter-particular aggregation by the steric effect of the flexible PEG 

chains. 23,24 Depending on the length (molecular weight) and density (the number of PEG present on 

the surface), the PEG chains around the PEC are often enough to shield the surface charge of the 

nonstoichiometric complexes to achieve charge neutrality of the complex surface. 25–28 

Block or graft copolymers consisting of a cationic polymer and PEG were reported to form PEC 

micelles by interacting with negatively charged ODN or plasmid DNA. When a poly(1-lysine)-PEG 

(PLL-b-PEG) di-block copolymer interacts with negatively charged ODN or plasmid DNA, selfassembled 

micelles can generate. 18 The charge neutralization between polycation and negatively 

charged polynucleotide could form a core-shell type nanoparticulate, entrapped and surrounded by 

a hydrophilic PEG segment. Similarly, the formation of self-associated micelles based on the 

interaction between counter ions could also be observed in a mixture of PEG-grafted PEI and 

alkyl sulfates. 19 The surrounded hydrophilic shell not only stabilizes the micelles, but also increase 

water solubility of the particles. The PEC micelles formed from PLL-b-PEG, and ODN showed a 

spherical shape with a diameter of about 50 nm as determined by dynamic light scattering. The 

formation of PEC micelles also greatly enhanced the stability of particles. Overnight incubation at 

room temperature and even storage under frozen conditions did not affect the size distribution of the 

micelles. 18 The addition of oppositely charged poly(aspartic acid) to PLL-b-PEG/DNA PEC 

micelles caused exchange reaction between poly(aspartic acid) and DNA, leading to a release of 

DNA in the medium. 29 This result suggested that the PEC micelle-based delivery system could 

stabilize the complex and increase its water solubility by efficiently compartmentalizing it within 

PEG shells while retaining its DNA releasing capability by exchange reaction with counter ions. 

The PEC micelles also efficiently protect the encapsulated ODN or plasmid DNA from enzymatic 

attack. 30,31 Environmental stimuli-sensitive PEC micelles were prepared using thiolated PLL-b- 

PEG. The micelles formed from thiolated PLL-b-PEG and ODN were dissociated to release ODN 

in the presence of reduced glutathione, suggesting that the ODN entrapped in PEC micelles could 

be released in a cytoplasmic environment for proper therapeutic activity of ODN. 32 Other polyamine-PEG 

conjugates including PEI–PEG and polyspermine (PSP)-PEG also showed PEC 

micelles forming capability when introduced to ODN. 33,34 Diameters of the PEC micelles from 

PEI–PEG and PSP–PEG were ca. 32 nm and ca. 12 nm as determined by dynamic light scattering 

and transmission electron microscopy (TEM). The micelles could be stably stored in a solution for 

several months. The micellar formulation could efficiently minimize the nonspecific interaction 

with a major component of serum proteins, namely, bovine serum albumin (BSA). 

Recently, an alternative class of PEC micelles-based ODN delivery systems where PEC 

micelles could form from the interaction between PEG-conjugated oligonucleotide and polycation 

was introduced. Figure 24.1 shows the schematic diagram of the formation of PEC micelles. An 

A–B block-type conjugate, ODN–PEG, was synthesized. 20 An acid labile phosphoroamidate 

linkage was employed between ODN and PEG as a stimuli-sensitive linkage for the efficient 

release of ODN from the micelles in acidic endosomal pH. The ODN–PEG conjugate could 

self-associate to form PEC micelles, by interacting with several functional polycations such as 

fusogenic KALA peptide, PEI, and protamine (Figure 24.2). The charge-neutralized 


480 

O 

HN 

O 

O O O N O 

H H O 

H 

H OH 

ODN 

O 

HN 

O 

O O O N O 

H 

H 

O 

H OH 

H 

N 

O 

PEG 

Cationic polymer 

FIGURE 24.1 The schematic diagram of the formation of PEC micelles. 

Protective PEG 

shell 

Polyelectrolyte 

complex core 

ODN/polycation complex was entrapped and compartmentalized by surrounding PEG corona to 

form a stable core-shell structure. The spontaneously formed PEC micelles showed a spherical 

shape with a diameter of ca. 70 nm and a narrow distribution. 35 The PEC micellar formulation 

demonstrated much higher intracellular transport efficiency when compared to naked ODN. When 

formulated with antisense ODN corresponding to c-myc or c-raf mRNA, the PEC micelles could 

remarkably inhibit the proliferation of smooth muscle cells (A7R5) or ovarian cancer cells (A2780), 

respectively. 35,36 The formation of PEC micelles between ODN–PEG conjugate and linear PEI 

provides ODN with enhanced nuclease resistance. 37 The PEC formulation also showed minimal 

adsorption of serum proteins, suggesting that the surrounding PEG shell could provide the nanoparticulate 

with improved water solubility and with an efficient protection from the nonspecific or 

Percent cell growth (%) 

100 

80 

60 

40 

0 

KALA 

PEI Protamine no ODN 


FIGURE 24.2 Influence of PEC micelles containing antisense c-myb on the proliferation of smooth muscle 

cells. Each PEC micelle was prepared by ODN–PEG conjugate with different polycations, KALA, PEI, and 

protamine. The size and surface zeta-potential values of the different micellar formulations were similar to 

those of the ODN–PEG/KALA formulation. The concentration of antisense ODN used in each formulation 

was 20 mg/ml. (From Jeong, J. H., Kim, S. W., and Park, T. G., Bioconjug. Chem., 14, 2003. With permission.) 



specific interaction with extracellular or intracellular proteins. These features would make PEC 

micellar formulation feasible for in vivo applications, especially for a systemic administration. 

24.2.3 IN VIVO APPLICATIONS 

There are two important factors that should be considered in designing polymeric gene carriers for 

in vivo application: pharmacokinetic behaviors in systemic circulation and cellular uptake and 

intracellular trafficking of the carriers. Systemically administered polymeric gene carriers could 

experience a number of barriers upon injection, including the attacks of extracellular nucleases and 

systemic clearance systems such as RES. The accumulation of the carriers in a desired tissue or 

organ could also be one of the barriers the carriers should overcome. Even after arriving at the 

appropriate destination, there are more cellular barriers, including cellular uptake, escape from 

endosomal compartment, and localization into the nucleus when it is required. A number of 

previous reports on micelle-based drug delivery systems demonstrated improved pharmacokinetics 

and cellular internalization. 21,38–42 The PEC micelles have a similar core-shell structure to polymeric 

micelles. The surrounding hydrophilic shell could minimize the nonspecific adsorption of 

serum proteins to avoid the surveillance of the systemic clearance systems, allowing prolonged 

circulation of the carrier. Also, the colloidal size of PEC micelles makes it feasible for them to be 

used for passive tumor targeting. Nanoparticulates, generally having less than 100 nm, could be 

expected to experience enhanced permeability and retention in the vicinity of a solid tumor. 43 In 

addition, the use of functional polymers that can facilitate the endosomal escape of ODN will be 

advantageous to improve intracellular trafficking. For example, PEI, one of the commonly used 

cationic gene carriers, has a high buffering capacity that may facilitate escape of the complexes 

from the endosomal compartment. 44 

In vivo anti-tumor activity of the PEC micellar formulation (ODN–PEG/PEI) containing antisense 

ODN that blocks the expression of c-raf gene was evaluated by intratumoral injection in nude 

mice with subcutaneously implanted solid tumor xenograft derived from human ovarian cancer. 18 

Significant retardation of solid tumor growth was observed in PEC formulation with antisense c-raf 

ODN compared to the formulations with a scrambled ODN and PBS (Figure 24.3). When the PEC 

micelles containing antisense c-raf ODN were intravenously administered to the mice bearing a 

Tumor volume (cm 3 ) 

8 

6 

4 

2 

0 

0 2 

PEG miceless (antisense ODN) 

PEG miceless (mismatched ODN) 

Antisense ODN alone 

Control 

4 6 

8 10 12 14 16 18 20 

Time (Day) 

FIGURE 24.3 Effect of ODN–PEG/PEI PEC micelles containing antisense c-raf ODN on the growth of 

human ovarian cancer-derived tumor xenograft in nude mice. Each ODN formulation was administered 

intratumorally at 2.5 mg kg K1 injection K1 . (From Jeong, J. H., Kim, S. W., and Park, T. G., J. Control. 

Rel., 93, 2003. With permission.) 


482 


2000 

1500 

1000 

500 

Antisense MC 

Mismatched MC 

Control 

Antisense ODN Only 

0 

0 2 4 6 

Time (days) 

solid tumor derived from human lung cancer, the proliferation rate of the tumor was significantly 

reduced (Figure 24.4). 35 The PEC micellar formulation also exhibited much higher deposition of 

ODN in a solid tumor region than in a naked ODN after 12 h of intravenous injection through a tail 

vein (Figure 24.5). Taken together, the results suggested the hydrophilic PEG shell surrounding 

charge-neutralized polyelectrolyte core could successfully inhibit the nonspecific interaction of 

serum proteins in the blood stream. Also, through loosened vasculature near the rapidly growing 

tumors, the nanosized particles could be targeted to the solid tumor by the enhanced permeation and 

retention (EPR) effect. 43 

8 

10 12 14 16 

FIGURE 24.4 Effect of systemically administered ODN–PEG/PEI PEC micelles containing antisense c-raf 

ODN on the growth of human lung cancer-derived tumor xenograft in nude mice. Each ODN formulation was 

administered through tail vein at 2.5 mg kg K1 injection K1 . (From Jeong, J. H., Kim, S. H., Kim, S. W., and 

Park, T. G., Bioconjug. Chem., 16(4), 2005. With permission.) 

% ID/g tissue 

1.4 

1.2 

1.0 

0.8 

0.6 

0.4 

0.2 

0.0 

3 h 

12 h 

24 h 

ODN PEC Micelles 


FIGURE 24.5 Accumulation of ODN–PEG/PEI PEC micelles in solid tumor (human lung cancer) after 

intravenous administration. (From Jeong, J. H., Kim, S. H., Kim, S. W., and Park, T. G., Bioconjug. Chem., 

16(4), 2005. With permission.) 



24.3 LIGAND-MEDIATED TARGETING OF ANTISENSE OLIGONUCLEOTIDES 

Nonviral carriers do not have an inherent selectivity of cell surface binding and internalization. The 

limitation can be addressed by conjugating various ligands targeting specific cells or tissure to 

polymeric gene carriers. The PEC micelles demonstrated promising potential as a carrier for 

systemic cancer treatment that could target solid tumors. In conjunction with the passive targeting 

property, it would be advantageous if the PEC micelles have cell- or tissue-specific ligands on the 

surface. The specific ligand could play an additional role for enhanced internalization of PEC 

micelles into the cells by receptor mediated endocytosis. 

24.3.1 LIGAND-MEDIATED ODN TARGETING SYSTEMS BASED ON PEC MICELLES 

A graft copolymer composed of PEI and PEG where the terminal end of PEG was biotinylated for 

further conjugation was synthesized and used for ligand-mediated targeting of cancer cells. PEC 

micelles could be generated by interacting biotinylated PEI-g-PEG and ODN. Avidin was attached 

to the biotin located at the end of the PEG segment and biotinylated trasferrin (Tf) was subsequently 

attached to the avidin to obtain a PEI–PEG-Tf conjugate. 45 Enhanced cellular uptake of PEI–PEG- 

Tf/ODN PEC micelles was shown in cancer cells. When formulated with antisense ODN targeting 

multi-drug efflux transporter protein, P-glycoprotein (P-gp), the PEI–PEG-Tf/ODN PEC micelles 

efficiently inhibit the expression of the P-gp, leading to intracellular accumulation of a P-gp specific 

probe (rhodamine 123) in multi-drug resistant (MDR) cancer cells. 

Based on the same idea of PEG–ODN, a small interference lactosylated PEG-siRNA (Lac- 

PEG-siRNA) conjugate was recently synthesized. 46 An acid-labile b-thiopropionate linkage was 

introduced between siRNA and PEG to facilitate acid-triggered intracellular release of siRNA in 

the cytoplasm. The PEC micelles were prepared by interacting the PEG-siRNA conjugate with 

PLL, and they formed spherical shape nanoparticulates with an average diameter of. 117 nm. The 

lactose moiety at the end of the PEG segment could target the asialogycoprotein of hepatocytes to 

facilitate the receptor-mediated endocytosis of the PEC micelles. The PEC micelles demonstrated a 

significant RNA interference (RNAi) activity in human hepatoma cells (HuH-7), comparable to that 

of commercial cationic liposome-based carrier oligofectAMINE. 

24.3.2 IN VIVO APPLICATIONS 

PEC micelles demonstrated potential as a tumor targeting carrier for antisense ODN. Owing to their 

small size (generally less than 100 nm), the PEC micelles could be passively located in a solid 

tumor region through the loosened endothelial vasculature of the tumors by EPR effect. In addition 

to the passive targeting property, it would be more desirable if the carrier could also target a specific 

cell or tissue. Folate (FA) has been recognized as a potential targeting ligand for cancer cells. 

Expression of FA receptor is frequently up-regulated in a number of malignant tumors and highly 

restricted in normal cells and tissues. 47,48 Therefore, FA was coupled to several polymeric carriers 

for drug targeting to cancer cell. 38,49,50 FA has also been used for the modification of cationic 

polymers for tumor-targeted gene delivery. 51–53 Recently, ODN–PEG conjugate having FA at the 

end of PEG was synthesized for the systemic treatment of cancer based on PEC micelle delivery 

system. 54 The size of ODN–PEG–FA/PEI PEC micelles was ca. 90 nm with narrow distribution. 

Systemic administration of the micellar formulation to nude mice bearing human tumor xenograft 

derived from KB cells, FA-over-expressed the cells (KB cells), demonstrating the enhanced deposition 

of ODN in solid tumor region when compared to ODN–PEG/PEI PEC micelles. However, no 

significance in tumor deposition of ODN was observed between ODN–PEG–FA and ODN–PEG 

formulations in FA-receptor negative tumor xenograft model (A549 cells). 


484 


Polymeric gene carriers have been developed as substitutes for viral vectors that could cause a 

series of unexpected side effects such as immunogenecity and oncogenecity. Micelle-based gene 

carriers exhibited a number of advantageous features for systemic delivery, including colloidal 

stability in solution, reduced interaction with components of biological fluid, improved in vivo 

pharmacokinetics, and the possibility of passive and active targeting of cancers. The intracellular 

pharmacokinetic behavior of the nucleic acid-based therapeutics could also be improved by 

employing functional polymers with the existing formulation. Although there is still room for 

further improvement, the micellar gene delivery system could be considered a good example of 

a multi-functional drug delivery platform that has the potential to overcome a number of intrinsic 

barriers in the in vivo application of therapeutic genes. 

REFERENCES 


1. Simons, M., Edelman, E. R., DeKeyser, J. L., Langer, R., and Rosenberg, R. D., Antisense c-myb 

oligonucleotides inhibit intimal arterial smooth muscle cell accumulation in vivo, Nature, 359, 67, 

1992. 

2. Stein, C. A. and Cheng, Y. C., Antisense oligonucleotides as therapeutic agents—is the bullet really 

magical?, Science, 261, 1004, 1993. 

3. Elbashir, S. M., Harborth, J., Lendeckel, W., Yalcin, A., Weber, K., and Tuschl, T., Duplexes of 

21-nucleotide RNAs mediate RNA interference in cultured mammalian cells, Nature, 411, 494, 2001. 

4. Elbashir, S. M., Lendeckel, W., and Tuschl, T., RNA interference is mediated by 21- and 22-nucleotide 

RNAs, Genes Dev., 15, 188, 2001. 

5. Wagner, R. W., Gene inhibition using antisense oligodeoxynucleotides, Nature, 372, 333, 1994. 

6. Jeong, J. H. and Park, T. G., Novel polymer-DNA hybrid polymeric micelles composed of hydrophobic 

poly(D,L-lactic-co-glycolic acid) and hydrophilic oligonucleotides, Bioconjug. Chem., 12, 917, 

2001. 

7. Braasch, D. A., Paroo, Z., Constantinescu, A., Ren, G., Oz, O. K., Mason, R. P., and Corey, D. R., 

Biodistribution of phosphodiester and phosphorothioate siRNA, Bioorg. Med. Chem. Lett., 14, 1139, 

2004. 

8. Agrawal, S., Goodchild, J., Civeira, M. P., Thornton, A. H., Sarin, P. S., and Zamecnik, P. C., 

Oligodeoxynucleoside phosphoramidates and phosphorothioates as inhibitors of human immunodeficiency 

virus, Proc. Natl. Acad. Sci. USA, 85, 7079, 1988. 

9. Nielsen, P. E., Egholm, M., Berg, R. H., and Buchardt, O., Peptide nucleic acids (PNAs): Potential 

antisense and anti-gene agents, Anticancer Drug Des., 8, 53, 1993. 

10. Egholm, M., Buchardt, O., Christensen, L., Behrens, C., Freier, S. M., Driver, D. A., Berg, R. H., Kim, 

S. K., Norden, B., and Nielsen, P. E., PNA hybridizes to complementary oligonucleotides obeying the 

Watson–Crick hydrogen-bonding rules, Nature, 365, 566, 1993. 

11. Pagnan, G., Stuart, D. D., Pastorino, F., Raffaghello, L., Montaldo, P. G., Allen, T. M., Calabretta, B., 

and Ponzoni, M., Delivery of c-myb antisense oligodeoxynucleotides to human neuroblastoma cells 

via disialoganglioside GD(2)-targeted immunoliposomes: Antitumor effects, J. Natl. Cancer Inst., 92, 

253, 2000. 

12. Juliano, R. L. and Akhtar, S., Liposomes as a drug delivery system for antisense oligonucleotides, 

Antisense Res. Dev., 2, 165, 1992. 

13. Barron, L. G., Meyer, K. B., and Szoka, F. C. Jr., Effects of complement depletion on the pharmacokinetics 

and gene delivery mediated by cationic lipid-DNA complexes, Hum. Gene Ther., 9, 315, 

1998. 

14. Meyer, O., Kirpotin, D., Hong, K., Sternberg, B., Park, J. W., Woodle, M. C., and Papahadjopoulos, 

D., Cationic liposomes coated with polyethylene glycol as carriers for oligonucleotides, J. Biol. 

Chem., 273, 15621, 1998. 

15. Kim, J. S., Kim, B. I., Maruyama, A., Akaike, T., and Kim, S. W., A new non-viral DNA delivery 

vector: The terplex system, J. Control. Rel., 53, 175, 1998. 



16. Hashida, M., Takemura, S., Nishikawa, M., and Takakura, Y., Targeted delivery of plasmid DNA 

complexed with galactosylated poly(L-lysine), J. Control. Rel., 53, 301, 1998. 

17. Nishikawa, M., Takemura, S., Takakura, Y., and Hashida, M., Targeted delivery of plasmid DNA to 

hepatocytes in vivo: Optimization of the pharmacokinetics of plasmid DNA/galactosylated 

poly(L-lysine) complexes by controlling their physicochemical properties, J. Pharmacol. Exp. 

Ther., 287, 408, 1998. 

18. Kataoka, K., Togawa, H., Harada, A., Yasugi, K., Matsumoto, T., and Katayose, S., Spontaneous 

formation of polyion complex micelles with narrow distribution from antisense oligonucleotide and 

cationic block copolymer in physiological saline, Macromolecules, 29, 8556, 1996. 

19. Bronich, T. K., Cherry, T., Vinogradov, S. V., Eisenberg, A., Kabanov, V. A., and Kabanov, A. V., 

Self-assembly in mixtures of poly(ethyleneoxide)-graft-poly(ethylenimine) and alkyl sulfates, Langmuir, 

14, 6101, 1998. 

20. Jeong, J. H., Kim, S. W., and Park, T. G., Novel intracellular delivery system of antisense oligonucleotide 

by self-assembled hybrid micelles composed of DNA/PEG conjugate and cationic fusogenic 

peptide, Bioconjug. Chem., 14, 473, 2003. 

21. Yoo, H. S. and Park, T. G., Biodegradable polymeric micelles composed of doxorubicin conjugated 

PLGA–PEG block copolymer, J. Control. Rel., 70, 63, 2001. 

22. Jeong, J. H., Song, S. H., Lim, D. W., Lee, H., and Park, T. G., DNA transfection using linear 

poly(ethylenimine) prepared by controlled acid hydrolysis of poly(2-ethyl-2-oxazoline), J. Control. 

Rel., 73, 391, 2001. 

23. Lee, H., Jeong, J. H., and Park, T. G., PEG grafted polylysine with fusogenic peptide for gene 

delivery: High transfection efficiency with low cytotoxicity, J. Control. Rel., 79, 283, 2002. 

24. Lee, H., Jeong, J. H., and Park, T. G., A new gene delivery formulation of polyethylenimine/DNA 

complexes coated with PEG conjugated fusogenic peptide, J. Control. Rel., 76, 183, 2001. 

25. Brus, C., Petersen, H., Aigner, A., Czubayko, F., and Kissel, T., Efficiency of polyethylenimines and 

polyethylenimine-graft-poly(ethylene glycol) block copolymers to protect oligonucleotides against 

enzymatic degradation, Eur. J. Pharm. Biopharm., 57, 427, 2004. 

26. Brus, C., Petersen, H., Aigner, A., Czubayko, F., and Kissel, T., Physicochemical and biological 

characterization of polyethylenimine-graft-poly(ethylene glycol) block copolymers as a delivery 

system for oligonucleotides and ribozymes, Bioconjug. Chem., 15, 677, 2004. 

27. Kunath, K., von Harpe, A., Petersen, H., Fischer, D., Voigt, K., Kissel, T., and Bickel, U., The 

structure of PEG-modified poly(ethylene imines) influences biodistribution and pharmacokinetics 

of their complexes with NF-kappaB decoy in mice, Pharm. Res., 19, 810, 2002. 

28. Jeong, J. H., Lee, M., Kim, W. J., Yockman, J. W., Park, T. G., Kim, Y. H., and Kim, S. W., Anti-GAD 

antibody targeted non-viral gene delivery to islet beta cells, J. Control. Rel., 107, 562, 2005. 

29. Katayose, S. and Kataoka, K., Water-soluble polyion complex associates of DNA and poly(ethylene 

glycol)-poly(1-lysine) block copolymer, Bioconjug. Chem., 8, 702, 1997. 


supramolecular assembly with poly(ethylene glycol)-poly(L-lysine) block copolymer, J. Pharm. Sci., 

87, 160, 1998. 

31. Harada, A., Togawa, H., and Kataoka, K., Physicochemical properties and nuclease resistance of 

antisense–oligodeoxynucleotides entrapped in the core of polyion complex micelles composed of 

poly(ethylene glycol)-poly(L-lysine) block copolymers, Eur. J. Pharm. Sci., 13, 35, 2001. 

32. Kakizawa, Y., Harada, A., and Kataoka, K., Glutathione-sensitive stabilization of block copolymer 

micelles composed of antisense DNA and thiolated poly(ethylene glycol)-block-poly(L-lysine): A 

potential carrier for systemic delivery of antisense DNA, Biomacromolecules, 2, 491, 2001. 

33. Vinogradov, S. V., Bronich, T. K., and Kabanov, A. V., Self-assembly of polyamine-poly(ethylene 

glycol) copolymers with phosphorothioate oligonucleotides, Bioconjug. Chem., 9, 805, 1998. 

34. Kabanov, V. A. and Kabanov, A. V., Interpolyelectrolyte and block ionomer complexes for gene 

delivery: Physico-chemical aspects, Adv. Drug Deliv. Rev., 30, 49, 1998. 

35. Jeong, J. H., Kim, S. H., Kim, S. W., and Park, T. G., Polyelectrolyte complex micelles composed of 

c-raf anti-sense oligodeoxynucleotide-poly(ethylene glycol) conjugate and poly(ethylenimine): Effect 

of systemic administration on tumor growth, Bioconjug. Chem., 16(4), 1034–1037, 2005. 

36. Jeong, J. H., Kim, S. W., and Park, T. G., A new antisense oligonucleotide delivery system based on 

self-assembled ODN–PEG hybrid conjugate micelles, J. Control. Rel., 93, 183, 2003. 


486 


37. Oishi, M., Sasaki, S., Nagasaki, Y., and Kataoka, K., pH-responsive oligodeoxynucleotide (ODN)poly(ethylene 

glycol) conjugate through acid-labile beta-thiopropionate linkage: Preparation and 

polyion complex micelle formation, Biomacromolecules, 4, 1426, 2003. 


J. Control. Rel., 96, 273, 2004. 

39. Yoo, H. S., Lee, E. A., and Park, T. G, Doxorubicin-conjugated biodegradable polymeric micelles 

having acid-cleavable linkages, J. Control. Rel., 82, 17, 2002. 

40. Harada-Shiba, M., Yamauchi, K., Harada, A., Takamisawa, I., Shimokado, K., and Kataoka, K., 

Polyion complex micelles as vectors in gene therapy–pharmacokinetics and in vivo gene transfer, 

Gene Ther., 9, 407, 2002. 


compounds, Adv. Drug Deliv. Rev., 54, 203, 2002. 

42. Ideta, R., Yanagi, Y., Tamaki, Y., Tasaka, F., Harada, A., and Kataoka, K., Effective accumulation of 

polyion complex micelle to experimental choroidal neovascularization in rats, FEBS Lett., 557, 21, 

2004. 


effect in macromolecular therapeutics: A review, J. Control. Rel., 65, 271, 2000. 

44. Boussif, O., Lezoualc’h, F., Zanta, M. A., Mergny, M. D., Scherman, D., Demeneix, B., and Behr, 

J. P., A versatile vector for gene and oligonucleotide transfer into cells in culture and in vivo: Polyethylenimine, 

Proc. Natl. Acad. Sci. USA, 92, 7297, 1995. 


corona for receptor-mediated delivery of oligonucleotides into cells, Bioconjug. Chem., 

10, 851, 1999. 

46. Oishi, M., Nagasaki, Y., Itaka, K., Nishiyama, N., and Kataoka, K., Lactosylated poly(ethylene 

glycol)-siRNA conjugate through acid-labile beta-thiopropionate linkage to construct pH-sensitive 

polyion complex micelles achieving enhanced gene silencing in hepatoma cells, J. Am. Chem. Soc., 

127, 1624, 2005. 

47. Lu, Y. and Low, P. S., Folate-mediated delivery of macromolecular anticancer therapeutic agents, 


48. Antony, A. C., The biological chemistry of folate receptors, Blood, 79, 2807, 1992. 

49. Kim, S. H., Jeong, J. H., Chun, K. W., and Park, T.G, Target-specific cellular uptake of PLGA 

nanoparticles coated with poly(1-lysine)-poly(ethylene glycol)-folate conjugate, Langmuir, 21, 

8852, 2005. 

50. Yoo, H. S. and Park, T. G., Folate-receptor-targeted delivery of doxorubicin nano-aggregates stabilized 

by doxorubicin-PEG-folate conjugate, J. Control. Rel., 100, 247, 2004. 

51. Cho, K. C., Kim, S. H., Jeong, J. H., and Park, T. G., Folate receptor-mediated gene delivery using 

folate-poly(ethylene glycol)-poly(L-lysine) conjugate, Macromol. Biosci., 5, 512, 2005. 

52. Benns, J. M., Maheshwari, A., Furgeson, D. Y., Mahato, R. I., and Kim, S. W., Folate-PEG-folategraft-polyethylenimine-based 

gene delivery, J. Drug Target., 9, 123, 2001. 

53. Kim, S. H., Jeong, J. H., Cho, K. C., Kim, S. W., and Park, T. G., Target-specific gene silencing by 

siRNA plasmid DNA complexed with folate-modified poly(ethylenimine), J. Control. Rel., 104, 223, 

2005. 

54. Jeong, J. H., Kim, S. H., Kim, S. W., and Park, T. G., In vivo tumor targeting of ODN–PEG-folic 

acid/PEI polyelectrolyte complex micelles, J. Biomater. Sci. Polym. Ed., 16, 1409, 2005. 


25 

CONTENTS 

Dendrimers as Drug and Gene 

Delivery Systems 

Tae-il Kim and Jong-Sang Park 

25.1 Introduction ....................................................................................................................... 489 

25.2 Drug Delivery .................................................................................................................... 490 

25.2.1 Introduction ......................................................................................................... 490 

25.2.2 Non-Covalent Encapsulation of Drugs ............................................................... 490 

25.2.3 Covalent Conjugation of Drugs .......................................................................... 493 

25.2.4 Neutron Capture Therapy.................................................................................... 494 

25.2.4.1 Boron Neutron Capture Therapy ....................................................... 494 

25.2.4.2 Gadolinium Neutron Capture Therapy .............................................. 495 

25.2.5 Photodynamic Therapy ....................................................................................... 496 

25.3 Gene Delivery.................................................................................................................... 498 

25.3.1 Introduction ......................................................................................................... 498 

25.3.2 PAMAM Dendrimers .......................................................................................... 498 

25.3.3 Polypropylenimine Dendrimers .......................................................................... 501 

25.3.4 Other Dendrimers ................................................................................................ 502 

25.4 Conclusion ......................................................................................................................... 504 


References..................................................................................................................................... 504 


Dendrimers are highly branched, mono-disperse macromolecules, characterized by layers from a 

core focal point called generations and multivalent end group functionalities. The word dendrimer 

originated from the Greek dendron, meaning “tree” and meros, meaning “part.” 

Dendrimers are prepared by iterative synthetic steps. First, Vögtle and coworkers reported the 

synthesis of dendritic structures as cascade molecules. 1 A lot of dendrimers have been developed, 

and they have become a subject of intense research, including within electrochemistry, photophysics, 

and supramolecular chemistry. 2,3 Over the past decade, characteristics such as a highly defined 

structure, modifiable surface functionality, and an internal cavity of dendrimers have aroused 

scientists’ interest in the biomedical field. In recent years, dendrimers have shown potential in 

fields, including drug delivery, gene delivery, magnetic resonance imaging (MRI), and anti-cancer 

therapeutics. 4–6 

The focus of this review is the history and the current trends of dendrimer application for drug 

and gene delivery systems. 


489

490 

25.2 DRUG DELIVERY 

25.2.1 INTRODUCTION 

The development of an efficient drug delivery system is very important to improve the pharmacological 

activity of drug molecules. Various types of drug delivery systems have been designed, 

including liposomes, polymeric micelles, etc. 7–9 Dendrimers have emerged as new alternatives and 

efficient tools for delivery of drug molecules. Figure 25.1 shows the molecular structures of various 

dendrimers used for delivery systems. Generally, bioactive drugs have hydrophobic moieties in 

their structure, resulting in low water-solubility that inhibits efficient delivery into cells. Dendrimers 

designed to be highly water-soluble and biocompatible have been shown to be able to improve 

drug properties such as solubility and plasma circulation time via various formulations and to 

deliver drugs efficiently. In comparison to linear polymeric carriers, the multivalent functionalities 

of dendrimers can be linked to drug molecules or ligands in a well-defined manner and can be used 

to increase the binding efficiency and affinity of therapeutic molecules to receptors via synergistic 

interaction. In addition, the low polydispersity of dendrimers can show a reproducible 

pharmacokinetic pattern. 

25.2.2 NON-COVALENT ENCAPSULATION OF DRUGS 

Drug delivery using unimolecular micelles has the advantage that the micellar structure is maintained 

and not disrupted at all concentrations, contrary to the behavior of conventional polymeric 

+ − − + 

− CO2 R4N 

+ NR4 − O2C 

+ − 

CO2 R4N NR4 O2C − + 

NR4 O2C 

CO2 R4N 

+ − − + 

NR4 O2C 

CO2 R4N 

+ − − + 

NR4 O2C 

CO2 R4N 

+ − − + 

NR4 O2C 

CO2 R4N 

+ NR4 − O2C 

+ − 

NR4 O2C 

+ − 

NR4 O2C 

CH3(OCH2CH2)16O 

CH3(OCH 2CH2)16O 

CH3(OCH 2CH2)16O 

CH 3(OCH2CH2)16O 

CH3(OCH2CH2)16O CH3(OCH2CH2)16O O(CH 2CH2O)16CH3 

O(CH 2CH2O)16CH3 

− + 

CO2 R4N 

CO2 − R4N + 

− + 

CO2 R4N 

+ − − + 

NR4 O2C 

CO2 R4N 

+ − − + 

NR4 O2C 

CO2 R4N 

+ − − + 

NR4 O2C 

CO2 R4N 

+ − − + 

NR4 O2C 

CO2 R4N 

+ − − + 

NR4 O2C 

CO2 R4N 

+ − − + 

NR4 O2C 

CO2 R4N 

+ − − + 

NR4 O2C − + − CO2 R4N 

+NR4− 

CO2 R4N NR4 O2C − + 

O2C 

CO2 R4N 

(a) (b) 

O 

O 

O 

O 

O O 

O O 

O 

O 

O 

O 

O 

O 

O 

O(CH2CH 2O)16CH3 

O(CH2CH2O)16CH3 O(CH2CH2O)16CH3 O(CH 2CH2O)16CH3 

(c) (d) 

O 

O 

O 

O 

O 

O 

NH2 H2N 

NH HN 

O 

O 

N 

H2N O 

HN 

HN 

O 

N 

NH 

HN 

O O N O 

HN 

H2N N 

H2N 

NH 

O 

NH2 

HN 

N O 

NH2 O 

O NH NH 

H 

N N 

N 

O O NH 

NH 

O 

NH2 N 

HN 

H2N O 

O 

O 

HN 

N 

N 

N 

H 

HN O 

HN 

O 

H2N O N 

NH O 

H2N NH 

NH2 NH 

O 

NH2 N O O 

NH 

HN 

O N 

NH 

NH 

O 

N 

NH2 O O 

NH NH 

H2N NH2 HO 

HO 

HO 

HO 

HO 

HO 

O 

O 

O 

O 

HO 

HO 

O 

O 

O 

O 

HO 

O 

O 

O 

O 

HO 

O 

O 

HO HO 

OH 

OH 

O O 

OH 

O O 

OH 

O 

O 

O 

O O O O 

O 

O 

O O 

FIGURE 25.1 Structures of various dendrimers. (a) Unimolecular micelle, (b) poly(amidoamine) dendrimer, 

(c) polyaryl ether dendrimer, (d) polyester dendrimer based on 2,2-bis(hydroxymethyl)propionic acid. 



O 

O 

O 

O 

O 

O 

O 

O 

O 

OH 

OH 

O 

O 

O 

O 

OH 

OH 

OH 

OH 

OH 

OH

Dendrimers as Drug and Gene Delivery Systems 491 

micelles because all segments are covalently connected. Newkome et al. reported that a dendritic 

unimolecular micelle was shown to solubilize various hydrophobic guest molecules in the internal 

hydrophobic cavities. 10 However, the problem of these systems for drug delivery is their high 

hydrophobicity. So, water-soluble and biocompatible poly(ethylene glycol) (PEG) has been introduced 

to the drug delivery systems. 

Initially, dendrimers were used for delivery systems mediated by non-covalent encapsulation of 

drugs as unimolecular micelles and dendritic boxes. Meijer and coworkers first reported the 

dendritic box based on the construction of a chiral shell of protected amino acids onto poly(propyleneimine) 

(PPI) dendrimers with 64 amine end groups (Figure 25.2). 11 They showed that the 

guest’s and cavity’s shapes determine the number of guests entrapped in the dendritic box; the 

architecture of the dendrimer is important, and shape-selective liberation of guests could be 

achieved by removing the shell in two steps. 12 Meijer and coworkers also developed a new 

host–guest system by synthesizing oligoethyleneoxy-modified PPI dendrimers, and they studied 

the interactions in aqueous media by UV/Vis titration. 13 PEGylated PPI dendrimers were also 

prepared as drug delivery systems by Pales and coworkers. 14 PEGylated PPI dendrimers were 

O 

O O 

o 

N 

N 

H 

N 

N 

N 

H 

N 

N 

H 

N N 

H 

N 

N 

N 

H 

H 

N 

N 

N 

H 

N 

H 

N N 

N 

N 

H 

N 

N 

N 

H 

N 

N 

R O 

O 

O 

O O 

O 

N 

O 

O 

N 

O O 

N 

N 

O 

N 

O 

N 

N O 

N N 

N O O 

O 

N N 

N 

O N 

N 

N O 

N 

O 

N 

N 

N 

N 

N 

N 

N 

N 

O 

O O O 

O O 

NH HN 

O O 

NH HN 

O NH 

O 

O 

HN NH 

NH HN 

O 

O 

O 

NH 

HN 

O 

HN 

O O 

O HN 

O 

O 

O 

O 

O 

HN 

HN 

NH 

HN 

NH HN 

HN NH 

NH 

NH HN 

O O 

HN 

HN NH 

NH HN 

HN 

NH 

NH 

H 

NH 

H 

H 

H 

H 

H 

H 

H 

H 

NH 

H 

NH 

NH 

HN 

NH HN 

HN NH 

HN 

NH HN 

O 

O 

O 

O 

O 

O 

O O 

O 

O 

O 

O O 

O 

O 

O 

OO 

O O 

O 

O O 

O 

R O R 

O 

R 

R 

R 

R 

RO 

R 

R 

RO 

R 

R 

R 

R 

R R 

R 

R 

R 

R 

R 

R 

R 

R 

R 

O 

O 

R 

O 

O O 

R 

O 

O 

R 

O O 

O 

R O 

O 

R OO R O 

O O 

O O 

Guest molecule 

FIGURE 25.2 Scheme of modified PPI dendrimers for the non-covalent encapsulation of guest molecules. 


R

492 


shown to enhance water-solubility of encapsulated pyrene and betamethasones, securing their 

application as promising controlled release drug carriers. Encapsulated molecules were found to 

be solubilized in both regions of the interior of dendrimers and the PEG coat. The same group 

reported that the introduction of quaternary ammonium groups to PPI dendrimers not only 

enhanced the water solubility, but they also affected the protonation titration profile, leading to 

the release of the entrapped molecule within a narrower pH region. 15 Pyrene is released when the 

internal tertiary amines get protonated between pH 4 and 2, suggesting these dendrimers are 

potential candidates for pH-sensitive controlled-release drug delivery systems. 

Poly(amidoamine) (PAMAM) dendrimer that was first developed by Tomalia et al. 16 also has 

been extensively studied for non-covalent encapsulation of drug molecules. PAMAM dendrimers 

are water-soluble and biocompatible, and they possess internal cavities hosting guest molecules. 

Twyman et al. modified the ester end groups of half-generation PAMAM dendrimers with Tris. 17 

This highly water-soluble dendrimer could solubilize hydrophobic guest molecules at pH 7.0, but it 

lost its binding capability to guest molecules at pH 2.0, presumably because of the protonation of 

the internal tertiary amines. 

To improve biocompatibility and water-solubility, PEG was grafted on the surface of PAMAM 

dendrimers. PEG-attached PAMAM dendrimers could encapsulate anti-cancer drugs, adriamycin, 

and methotrexate, and their ability to encapsulate these drugs increased with the increasing 

dendrimer generation and chain length of PEG grafts. 18 The dendrimers slowly released the 

drugs in low ionic strength solution; however, drugs were readily released in isotonic solutions. 

Another group explored the use of PEGylated PAMAM dendrimers for delivery of the anti-cancer 

drug 5-fluorouracil (5-FU) in vitro and in albino rats. 19 PEGylation of dendrimers was found to 

increase their drug-loading capacity and reduce the drug release rate and hemolytic toxicity, 

suggesting the use of PEGylated PAMAM dendrimers as prolonged delivery systems. 

Haba et al. prepared PEG-modified PAMAM dendrimers having methacryloyl groups on the 

chain ends of the dendrimer for drug delivery. 20 After linking of the methacryloyl groups, they 

found that the dendrimer decreased its affinity against rose Bengal because of its hiding the 

dendrimer interior of the peripheral network and that the dendrimer having the peripheral 

network retained Rose Bengal dye molecules tightly in their interior. This result indicated that 

the linking of polymerizable groups attached to the periphery of the dendrimer could be an efficient 

approach for retaining guest molecules in their interior, and that they could be used as photosensitizers 

for photodynamic therapy (PDT) and as boron10-containing compounds for neutron 

capture therapy because these pharmaceutical molecules could exhibit their therapeutic effects, 

even if they are not released from their carriers. 

Frechet and coworkers synthesized novel dendritic unimolecular micelles with a hydrophobic 

polyether dendrimer core surrounded by hydrophilic PEGs. 21 For the G-3 micelle, the loading of 

indomethacin was found to be of 11%, a value corresponding to approximately nine to ten drug 

molecules per micelle, and preliminary in vitro release tests showed that sustained release characteristics 

were achieved. Recently, linear dendritic block copolymers comprising PEG and either a 

polylysine or polyester dendrons were designed, and highly acid-sensitive cyclic acetals were 

attached to the dendrimer periphery for pH-responsive micelle systems. 22 At pH 7.4, the fluorescence 

of micelle-encapsulated Nile Red was constant, indicating it was retained in the micelle, 

whereas at pH 5, the fluorescence decreased, consistent with its release into aqueous solution. The 

rate of release was strongly correlated with the rate of acetal hydrolysis. These results suggest that 

the loading and controlled-release of drugs are dependent on the chemical structure of 

the dendrimer. 

A novel triblock copolymer composed of citric acid dendritic blocks and PEG core was 

synthesized and applied as a drug delivery system. 23 The dendrimers were found to be able to 

trap various hydrophobic drugs in their suitable sites and to release them in a controlled manner in 

in vitro experiments. 



25.2.3 COVALENT CONJUGATION OF DRUGS 

Another approach for dendritic drug carriers is based on covalent conjugation of drug molecules to 

the dendrimer periphery using the multivalency of dendrimers. Although many drug molecules 

were conjugated to the dendrimers through non-degradable linkages, introducing degradable bonds 

between the drug and the dendrimer can control the release of drug molecules. Figure 25.3 shows 

the scheme of drug-conjugated dendrimers. 

Dendrimers were first used for immunoconjugates. Chemical modification of the antibody with 

drug molecules often diminished its biological activity, but approaches to link the antibody to 

another macromolecule including the dendrimer were reported to increase drug loading while 

retaining the activity. Roberts et al. conjugated antibody to porphyrin to chelate radiolabeled 

copper ions using PAMAM dendrimer as a linker. 24 They found that 100% of the porphyrin 

bound to the heavy chain of the antibody, and dendrimer conjugates contributed 90% of the 

immunoactivity of unmodified antibody and showed a high radiolabeling level in contrast to the 

conjugates obtained through the random-coupling method. PAMAM dendrimers were also 

modified chemically by reaction with 1,4,7,l0-tetraazacyclododecanetetraacetic acid (DOTA) 

and diethylenetriaminepentaacetic acid (DTPA) type bifunctional metal chelators and used as a 

linker to monoclonal antibody (moAb) 2E4. 25 The dendrimer–antibody conjugates were found to 

retain biological activity and to be easily radiolabeled, suggesting use of these dendrimer conjugates 

for moAb-based radiotherapy or imaging. Recently, Patri et al. conjugated anti-PSMA 

(prostate-specific membrane antigen) to PAMAM dendrimer labeled with FITC and capped with 

acetic anhydride to minimize the non-specific interaction with the cell surfaces for targeting 

prostate cancer. 26 This dendrimer conjugate showed great potential to target PSMA-positive 

LNCaP cell line with minimal loss of immunoreactivity. Meanwhile, PAMAM dendrimers were 

utilized to investigate the potential for cancer chemotherapy by conjugating with anti-cancer drugs. 

Cisplatin, a potent anti-cancer drug, was conjugated to PAMAM dendrimer, giving a dendrimer– 

platinate that was highly water-soluble and released platinum slowly in vitro. 27,28 This dendrimer 

conjugate showed more potent anti-tumor activity than cisplatin and selective accumulation in solid 

tumor tissue and was also less toxic than cisplatin. Zhuo et al. synthesized a water-soluble acetylated 

PAMAM dendrimer series with a core of 1,4,7,10-tetraazacyclododecane and linked it to 

1-bromoacetyl-5-fluorouracil to form dendrimer–5-FU conjugates. 29 Hydrolysis of the conjugates 

in a phosphate buffer solution could release free 5-FU in a generation-dependent manner, showing 

the dendrimer conjugate seems to be a promising carrier for the controlled release of anti-tumor 

drugs. Wang et al. used PAMAM dendrimers for doxorubicin (DOX) delivery. 30 They conjugated 

semitelechelic poly[N-(2-hydroxypropyl)-methacrylamide] macromolecules (ST-PHPMA,) to the 

periphery of PAMAM dendrimer and introduced DOX to the dendrimer conjugate, star-like HPMA 

copolymer. It showed lower cytotoxicity than other brush-like polymer-DOX in A2780 human 

ovarian carcinoma cell line and a slow rate of in vitro DOX release, demonstrating that the 

differences of DOX internalization because of the structure of the polymers would result in different 

Drug 

Drug 

Targeting molecule 

FIGURE 25.3 Schematic view of drug–dendrimer conjugate. A dotted line, degradable linkage; a black line, 

non-degradable linkage. 


494 

intracellular DOX concentrations with cytotoxicity changes. 5-Aminosalicylic acid (5-ASA) was 

also conjugated to PAMAM dendrimers using two different spacers containing azo-bond, p-aminobenzoic 

acid (PABA), and p-aminohippuric acid (PAH) for colonic delivery. 31 The dendrimer 

conjugates incubated in rat fecal contents showed colon-specific and prolonged release of 

5-ASA through reductive cleavage of azo-bond bycolonic bacteria, suggesting the potential for 

colon-specific drug delivery carriers. In addition, Quintana et al. attached folic acid to acetamide-, 

hydroxyl-, or carboxyl-capped PAMAM dendrimers conjugated with methotrexate (MTX) to target 

tumor cells through the folate receptor. 32 These dendrimer conjugates showed enhanced cellular 

uptake, and this targeted delivery improved the cytotoxic response of the cells to MTX 100-fold 

over the free drug. Recently, Tansey et al. also conjugated folic acid and indocyanine green, a 

model diagnostic agent to biodegradable poly(L-glutamic acid) (PG) chain-terminated PAMAM 

dendrimer for tumor targeting. 33 The PG–dendrimer conjugates showed a more prolonged 

degradation pattern than linear PG chain in the presence of lysosomal enzyme. After conjugation 

of PEG, the dendrimer conjugates were found to show reduced non-specific interaction and to bind 

selectively to tumor cells expressing folate receptors. 

Folic acids have already been conjugated to acid-labile hydrazide-terminated polyaryl ether 

dendrimers to target tumor cells by Frechet and coworkers. 34 The conjugates are soluble in aqueous 

medium above pH 7.4. The anti-tumor drug methotrexate was also attached to the dendrimers. The 

same group also prepared polyaryl ether dendrimer–PEG conjugates as model drug carriers. 35 

Cholesterol and two amino acid derivatives were attached to the dendrimer via carbonate, ester, 

and carbamate linkages. 

Greenwald and coworkers synthesized novel PEG–poly(aspartic) acid dendron conjugates and 

attached an anti-tumor agent, cytosine arabinoside (ara-C), to the dendron conjugates. 36 These 

conjugates showed greater tumor inhibition than free drugs and increased aqueous solubility, 

resulting in high loading and controlled release of drugs. Another study of conjugation of ara-C 

to PEG-dendrons composed of aminoadipic acid was accomplished by Schiavon et al. 37 The 

conjugates also presented prolonged blood residence time and higher loading of drugs. 

Polyester dendrimers based on the monomer unit 2,2-bis(hydroxymethyl)propanoic acid were 

also developed as drug carriers both in vitro and in vivo. 38 These dendrimers were found to be both 

water soluble and non-toxic. A 3-arm PEG–polyester dendrimer hybrid that conjugated with DOX 

via a hydrazone linkage susceptible to acidic hydrolysis showed pH-dependent drug release in vitro 

and a degree of anti-cancer activity in cancer cell lines. In addition, the little accumulation of the 

dendrimer conjugate found in any organ examined, including the liver, heart, and lungs, suggested 

that it is highly water soluble and biocompatible. 

Recently, a novel approach to developing covalently conjugated dendritic drug carriers was 

presented by two different groups. 39,40 These cascade-release or self-immolative conjugates were 

able to simultaneously release all drug molecules connected to the dendritic termini by a single 

trigger, resulting in whole dendrimer collapse. Moreover, Shabat and coworkers demonstrated selfimmolative 

dendrimers as a platform for multi-prodrugs, activated by a single enzymatic cleavage. 

41 The bioactivation of the heterodendritic prodrugs of DOX and campothecin (CPT) was 

evaluated using the Molt-3 leukemia cell line, and the conjugates showed a mild to significant 

increase in toxicity in comparison with the classical monomeric prodrugs. 

25.2.4 NEUTRON CAPTURE THERAPY 

25.2.4.1 Boron Neutron Capture Therapy 


Boron neutron capture therapy (BNCT) is a cancer-therapeutic application based on a nuclear 

capture reaction of a lethal 10 B(n,a) 7 Li 3C reaction. 42 If 10 B can be delivered at a concentration 

of at least 10 9 atoms per cell to tumor tissue, subsequent irradiation with thermal or epithermal 

neutrons produces highly energetic a particles and Li 3C ions that damage the tumor cells in nuclear 



fission reaction. To deliver high levels of boron in tumor tissue, it is proposed to use dendrimers as 

boron carriers for their well-defined molecular structure and multivalency. 

Initially, Barth et al. utilized PAMAM dendrimers in order to boronate moAb IB16-6 that 

targeted murine B16 melanoma using isocyanato polyhedral borane (Na(CH3)3NB10H8NCO). 43 

The boronated dendrimers were found to localize in the liver and spleen in in vivo distribution 

study, leaving a question of if the properties of boronated dendrimers can be modified to reduce 

their hepatic and splenic localization. Calssonn and coworkers attached epithermal growth factor 

(EGF) to boronated PAMAM dendrimers in order to target EGF receptor (EGFR) that is over 

expressed in brain tumors. 44 The boronated dendrimers-EGF were bound to malignant glioma cell 

membrane and endocytosed, resulting in the accumulation of boron in lysosome. Although these 

boronated dendrimers were localized in the liver and accumulated a little in the tumor after 

intravenous injection into rats having intracerebral C6 glioma transfected with EGFR gene, 

direct intratumoral injection could selectively deliver them to EGFR-positive gliomas. 45 Recently, 

use of another moAb that is directed against EGFR and EGFRvII, cetuximab (IMC-C225), was 

investigated for boronated PAMAM dendrimer-mediated BNCT. 46 There were w1100 boron 

atoms per molecule of cetuximab with only a slight reduction of Ka. Localization study of these 

conjugates in F98 rats bearing intracerebral implants of either EGFR-transfected or wild-type 

gliomas by intratumoral injection showed specific molecular targeting of EGFR and accumulation 

of boron. 

To reduce the hepatic uptake of boronated dendrimers, PEG was introduced to the dendrimers. 

47 Moreover, folic acid was also conjugated to the dendrimers to target the folate receptors in 

cancer cells. Among all the prepared combinations, boronated dendrimers with 1–1.5 PEG2000 

units exhibited the lowest hepatic uptake in C57BL/6 mice. Interestingly, dendrimers containing 

w13 decaborate clusters, w1 PEG2000 unit, and w1 PEG800 unit with folic acid attached to the 

distal end showed receptor-dependent uptake in contrast to dendrimers containing w15 decaborate 

clusters and w1 PEG2000 unit with folic acid attached to the distal end in in vitro studies using 

folate receptor (C) KB cells. Biodistribution studies with former dendrimer in C57BL/6 mice 

bearing folate receptor (C) murine 24JK-FBP sarcomas resulted in selective tumor uptake 

(6.0% ID/g tumor), and they also showed an increase in uptake of the dendrimers by the liver 

and kidney, warranting further effort in the optimization of biodistribution profiles of these 

boronated dendrimers. 

Polylysine dendrons also were examined for BNCT. Borane moieties were attached to the 

periphery of the dendrons and peptide spacers linked the dendrons to targeting molecules such 

as antibodies, a fluorescent moiety, and a PEG chain (Figure 25.4). 48 This offered a promising 

perspective for application in BNCT. 

25.2.4.2 Gadolinium Neutron Capture Therapy 

Although the vast majority of NCT research has involved various aspects of boron chemistry for 

BNCT, 157 Gd has also been suggested from theoretical and experimental studies with limited 

success. 49–51 To some extent, Gd complexes have been clinically employed as an MRI contrasting 

agent. 52 The potential to trace the in vivo pharmacokinetics of a Gd-NCT agent directly by MRI 

would be useful for planning the neutron irradiation and to predict the therapeutic effect. 

Kobayashi et al. synthesized avidin-G6-(1B4M-Gd)254(Av-G6Gd) from PAMAM dendrimer 

G6, biotin, avidin, and 2-(p-isothiocyanatobenzyl)-6-methyl-diethylenetriaminepentaacetic acid 

(1B4M). 53 A sufficient amount (162 ppm) of Av-G6Gd was accumulated and internalized into 

the SHIN3 cells both in vitro and in vivo to kill the cell using 157/155 Gd with external irradiation 

with an appropriate neutron beam while monitoring with MRI, suggesting that Av-G6Gd may be a 

promising agent for Gd neutron capture therapy of peritoneal carcinomatosis, and it may have the 

potential to permit monitoring of its pharmacokinetic progress with MRI. 


496 

H 2 N 

H 2 N 

N 

H 

O 

H N 2 

N 

H 

NH 

25.2.5 PHOTODYNAMIC THERAPY 

O 

O 

NH 

O 

NH 2 

HN 

O 

O 

O 

NH 

O O 

O 

O 

H 

H 

H 

N 

N 

N 

N 

N 

H 

H 

O O O 

HN 

NH 2 

HN 

O 

HN 

O 

NH 2 

O 

NH 

PDT is a technique for inducing tissue damage with light, following administration of a lightactivated 

photosensitizing drug that can be selectively concentrated in malignant or diseased 

tissue. 54 Activation of the photosensitizer results in the generation of reactive oxygen species 

(ROS), primarily singlet oxygen that oxidizes intracellular substrates such as lipids and amino 

acid residues, leading to cell death. The most widely studied application of PDT to date is in cancer 

therapy. However, the use of conventional photosensitizer systems bring skin phototoxicity, 

damage to normal tissue because of poor tumor selectivity, poor water solubility, and difficulties 

in treating solid tumors. In order to improve the disadvantages, dendrimers would be manufactured 

as promising photosensitizer carriers through modification of peripheral functionality. 

Battah et al. synthesized dendrimers with 5-aminolevulinic acid (ALA) moieties conjugated to 

the periphery by ester bonds that are cleavable by esterase inside the cells for PDT. 55 ALA is a 

natural precursor of an effective photosensitizer protophorphyrin IX (PpIX), and its administration 

can increased the cellular concentration of PpIX. 56 These ALA dendrimers showed increased 

NH 2 

N 

H 

NH 2 

NH 2 

NH 2 


O 

N 

H 

SH 

O 

HN SO 2 

O PEG 

FIGURE 25.4 Polylysine dendrimer conjugate used in boron neutron capture therapy (BNCT). 


N (CH 3 ) 2


production of PpIX and higher cytotoxicity after irradiation relative to free ALA in tumorigenic 

keratinocyte PAM 212 cell line. 

Another approach of designing dendritic photosensitizers for PDT involved aryl ether 

dendrimer porphyrins (DP) with either 32 quaternary ammonium groups or carboxylic acid 

moieties on their periphery (Figure 25.5). 57 Hydrophilic DPs were found to be localized in 

membrane-limited organelles but hydrophobic PIX diffused through the cytoplasm, except the 

nucleus. DPs showed remarkably higher 1 O2-induced cytotoxicity against Lewis lung carcinoma 

(LLC) cells and far lower dark toxicity than PIX, demonstrating their highly selective photosensitizing 

effect in combination with a reduced systemic toxicity. 

Zhang et al. entrapped DP with 32 amine groups on the periphery, [NH2CH2CH2NHCO]32DPZn 

with polyion micells composed of PEG and poly(aspartic) acid for effective PDT. 58 Compared to 

[NH2CH2CH2NHCO]32DPZn, a relatively low cellular uptake of [NH2CH2CH2NHCO]32DPZn 

X 

X 

X 

X 

X 

X 

X 

X 

X 

X 

O 

O 

O 

O 

O 

O 

O 

O 

O 

O 

X 

O 

O 

X 

O 

O 

X 

O O 

X 

O 

O 

O O 

O 

X X 

O O 

X 

X = CONH(CH 2) 2N + Me 3CI − or COO − H + 

FIGURE 25.5 Porphyrin-based dendrimer for photodynamic therapy. 


O 

O 

O 

O 

O 

O 

N 

N 

Zn 

N 

N 

O 

X 

O 

O 

O 

X 

O O 

O 

O O 

O 

O 

O 

O 

X 

O 

X 

O O 

O 

X 

O 

O 

O 

O 

X 

O 

O 

O 

O 

O 

O 

X 

X 

X 

X 

X 

X 

X 

X 

X

498 

incorporated in the PIC micelle was observed, yet the latter exhibited enhanced photodynamic 

efficacy on the LLC cell line. Also, this system was found to reduce the dark toxicity of the cationic 

dendrimer porphyrin, probably as a result of the biocompatible PEG shell of the micelles. 

Another approach for deeper tissue penetration was tried using two-photon absorption (TPA) 

with near-infrared lasers. Porphyrin sensitizers have low TPA cross-sections, limiting their usefulness 

in TPA applications. Recently, Dichtel et al. attached donor chromophores capable of efficient 

TPA to a central porphyrin acceptor for an enhanced TPA cross-section. 59 This dendrimer was 

found to enhance the effective TPA cross-section of a porphyrin acceptor, allowing more efficient 

generation of singlet oxygen using 780 nm light. This result showed a potential use of PDT to 

tumors below the skin surface. 

25.3 GENE DELIVERY 


Cationic dendrimers have been reported to deliver genetic materials into cells by forming 

complexes with negatively charged genetic materials through electrostatic interaction. The 

complexes bound to cell membranes were known to be internalized into cells by endocytosis 

(Figure 25.6). It has been postulated that the branched cationic polymers such as polyethyleneimine 

or PAMAM dendrimer, having a high buffer capacity lead to polymeric swelling in acidic endosomes, 

disrupting the membrane barrier and resulting in release of the complexes. 60 However, the 

uptake and the transfection mechanism of cationic polymers are not yet fully understood. Recently, 

two different mechanisms, caveolae and clathrin-mediated endocytosis, were also reported to 

involve the uptake mechanism of the polyplexes. 61 

Several dendrimers have been synthesized and used for gene delivery carriers from a decade 

ago. The history is not long, but the intense research on developing dendritic gene delivery carriers 

has been carried out because of many characteristic advantages of dendrimers such as monodisperse 

size of dendrimers and controllable multivalent end-group functionalities. 

25.3.2 PAMAM DENDRIMERS 

It is only for a decade that the potential of dendrimers as gene delivery carriers has been investigated. 

The first applied dendrimer for gene delivery was PAMAM dendrimer that was reported by 

Haensler and Szoka. 62 They showed that the dendrimers mediate high efficiency transfection of a 

variety of suspension- and adherent-cultured mammalian cells, and dendrimer-mediated 

DNA 

Dendrimer 

DNA-dendrimer 

complex 

Endocytosis 

FIGURE 25.6 Scheme of dendrimer-mediated transfection. 


Nucleus 

Endosome 


DNA 

transcription 

Protein 

translation


transfection is a function both of the dendrimer/DNA ratio and the diameter of the dendrimer. When 

GALA, a water soluble, membrane-destabilizing peptide, is covalently attached to the dendrimer 

via a disulfide linkage, transfection efficiency is increased by 2–3 orders of magnitude. Also, they 

concluded that the high transfection efficiency of the dendrimers may be caused by the low pKa of 

the amines in the polymer that permit the dendrimer to buffer the pH change in the endosomal 

compartment. Kukowska-Latallo et al. also reported a high degree of transfection of PAMAM 

dendrimers with minimal cytotoxicity in various cell lines. 63 The addition of DEAE-dextran 

increased the transfection efficiency, suggesting that dextran alters the nature of dendrimer/DNA 

complexes. The most efficient transfection was achieved with a lysosomotropic agent such as 

chloroquine, indicating that endosomal localization is a rate-determining step in transfection. 

Mazda and coworkers investigated the transfection of the Epstein–Barr virus (EBV)-based 

plasmid vector/PAMAM dendrimer complexes to improve the efficiency in hepatocellular carcinoma 

cells (HCC) and cholangiocarcinoma cells. 64,65 They found highly efficient suicide gene 

transfer into various HCC cells by EBV-based plasmid vectors/PAMAM dendrimer complexes 

in vitro, suggesting a possible application of this non-viral vector system to gene therapy of HCC, 

and they reported that EBV-based plasmid vectors/PAMAM dendrimer complexes equipped with 

the carcinoembryonic antigen (CEA) promoter showed high transfection in cholangiocarcinoma 

cells, providing an efficient non-viral method for the targeted gene therapy of CEAproducing 

malignancies. 

Thermally degraded PAMAM dendrimers with heterodisperse structure (fractured PAMAM) 

showed a dramatically enhanced transfection efficiency in in vitro experiments. 66 Transfection 

activity was found to be related both to the initial size of the dendrimer and to its degree of 

degradation. The increased transfection after the heating process was thought to be principally 

due to the increase in flexibility that enabled the fractured dendrimer to be compact when 

complexed with DNA and swell when released from DNA. The dendrimers are presently sold 

commercially by Qiagen as SuperFecte. 

Bielinska et al. investigated the ability of PAMAM dendrimers to function as an effective 

delivery system for antisense oligonucleotides (ODNs) and antisense expression plasmids for the 

targeted modulation of gene expression. 67 Transfections of ODNs or antisense cDNA plasmids 

usingPAMAMdendrimersresultedinaspecificand dose-dependent inhibition of luciferase 

expression. Binding of the phosphodiester oligonucleotides to dendrimers also extended their 

intracellular survival. Delong et al. also evaluated PAMAM dendrimer (generation 3, Mr 6909) 

as a potential delivery vehicle for oligonucleotides. 68 Use of dendrimers resulted in a 50-fold 

enhancement in cell uptake of oligonucleotide as determined by flow cytometry, enhanced cytosolic, 

and nuclear availability as shown by confocal microscopy. Interestingly, fluorescent dye, 

Oregon green 488-conjugated dendrimer, was proved to be a much better delivery agent for antisense 

compounds than unmodified dendrimer. 69 This result suggested that coupling of relatively 

hydrophobic small molecules to PAMAM dendrimers may provide a useful means of enhancing 

their capabilities as delivery agents for nucleic acids. 

The efficiency of cell transfection by PAMAM dendrimers was shown to be greatly increased 

when anionic oligomers (oligonucleotide or dextran sulfate) were mixed with plasmids before 

addition of the dendrimer. 70 This enhancement of gene transfer by anionic oligomers depended 

on their size, structure, and charge. This result may be due to modification of the interaction 

between plasmids and dendrimers induced by the oligonucleotide that effect leads to a less 

condensed structure of the polyplexes, thereby facilitating their penetration into cells. 

The transfection efficiency of PAMAM dendrimers was also increased by addition of substituted 

b-cyclodextrins (b-CD). 71 In vitro CAT expression was increased approximately 200-fold 

when dendrimer/DNA/b-CD formulations were applied on the surface of collagen membranes, 

indicating that b-CDs affected the physicochemical properties of dendrimer/DNA complexes. 

Besides, PAMAM dendrimers conjugated with a-cyclodextrins (a-CDs) were investigated for 

their potential as a gene delivery system. 72 The gene transfer activity of a-CD dendrimer conjugate 


500 


(G3) was higher than that of dendrimers (G2, G3, and G4) and of the a-CD conjugate (G2), and the 

cytotoxicity of the a-CD conjugate (G3) was lower than that of the a-CD conjugate (G4). This 

greatest gene transfer activity could be attributed to the synergistic interplay of the proton sponge 

effect of the dendrimer (G3) and the tentative membrane-disrupting effect of a-CD on endosomal 

membranes. Recently, the transfection efficiency of a-CD dendrimer conjugates was improved by 

surface modification of dendrimers with mannose. 73 The mannosylated conjugates were found to 

have a much higher gene transfer activity than unmodified a-CD dendrimer in various cell lines, 

which was independent of the expression of cell surface mannose receptors. Also, the conjugates 

provided higher gene transfer activity than dendrimer and a-CD dendrimer conjugate in kidney 

12 h after intravenous injection in mice. 

Arginine conjugation to the periphery of PAMAM dendrimer was reported to greatly increase 

the transfection efficiency. 74 Arginine residues of HIV-Tat peptide are thought to perform an 

important role in the intracellular translocation of molecular delivery. 75 The arginine-conjugated 

PAMAM dendrimer (PAMAM-Arg) showed enhanced gene expression in HepG2 cells, Neuro 2A 

cells, and primary rat vascular smooth muscle cells in comparison with native PAMAM dendrimer. 

It was presumed that the increased gene expression might be attributed to the difference in either the 

cell-penetrating activity during uptake, the nuclear localization efficiency after entry into the 

cytosol of arginine residues, or the synchronous function of both effects. Recently, Kang et al. 

conjugated Tat peptide to PAMAM dendrimer G5 for cellular delivery of antisense and siRNA 

oligonucleotides designed to inhibit MDR1 gene expression in vitro. 76 However, the dendrimer 

conjugate/oligonucleotide complexes weakly inhibited the gene expression, showing that conjugation 

of the dendrimer with the Tat cell-penetrating peptide failed to further enhance the 

effectiveness of the dendrimer. 

Lee et al. synthesized internally quaternized PAMAM-OH dendrimer (QPAMAM-OH) by 

methylation of internal tertiary amines and investigated the gene delivery potency. 77 Although 

the dendrimer did not show high transfection efficiency in in vitro experiments, it could form a 

complex with plasmid DNA unlike with unmodified PAMAM-OH dendrimer, and its cytotoxicty 

was revealed to be very low in comparison with amine-terminated PAMAM dendrimer. 

PAMAM dendrimers were also applied for ex vivo or in vivo gene delivery. Hudde et al. 

investigated the efficiency of activated polyamidoamine dendrimers (SuperFecte) to transfect 

rabbit and human corneas in ex vivo culture. 78 Whole-thickness rabbit or human corneas were 

transfected ex vivo with complexes of dendrimers and plasmids containing lacZ or TNFR-Ig genes. 

Six to ten percent of the corneal endothelial cells expressed the marker gene, and corneas transfected 

with TNFR-Ig plasmid were found to express TNFR-Ig protein. The fifth generation of intact 

PAMAM dendrimers was also investigated for its ability to enhance gene transfer and expression in 

a clinically relevant murine vascularized heart transplantation model. 79 The complexes of the 

plasmid pMP6A-beta-gal, encoding beta-galactosidase (beta-Gal), and dendrimer were perfused 

via the coronary arteries during donor graft harvesting. The grafts infused with pMP6A-beta-gal/ 

dendrimer complexes showed beta-Gal expression in myocytes from 7 to 14 days. The complexes 

containing 20 mg of DNA and 260 mg of dendrimer (1:20 charge ratio) in a total volume of 200 mL 

resulted in the highest gene expression in the grafts. Prolonged incubation (cold ischemic time) up 

to 2 h and pretreatment with serotonin were found to further enhance gene expression. 

Kukowska-Latallo et al. evaluated in vivo gene transfer and expression by intravascular and 

endobronchial routes using DNA complexed with G9 PAMAM dendrimer. 80 After intravenous 

administration of the dendrimer/pCF1CAT plasmid complexes, CAT expression was never 

observed in organs other than the lung. Repeated intravascular doses, administered four times at 

4-day intervals, maintained expression at 15–25% of peak concentrations achieved after the initial 

dose. However, in contrast with vascular delivery, intranasal delivery of the complex was found to 

give lower levels of CAT expression than that was observed with naked plasmid DNA. In situ 

localization of CAT enzymatic activity suggested that vascular administration seemed to achieve 

expression in the lung parenchyma, mainly within the alveoli, whereas endobronchial 



administration primarily targeted bronchial epithelium, showing that intravenously administered 

G9 dendrimer was an effective vector for pulmonary gene transfer and that transgene expression 

could be prolonged by repeated administration. PAMAM G4 dendrimer was also investigated for 

in vivo tumor targeting and transfection. 81 Biotinylated and 111 In labeled 20-mer antisense multiamino-linked 

oligonucleotide–avidin conjugate ( 111 In-oligo–Av) and the complexes of the 

oligonucleotide with PAMAM G4 ( 111 In-oligo/G4), biotinylated PAMAM G4–avidin 

( 111 In-oligo/G4–Av) were found to significantly enhance the tumor delivery of 111 In-oligo 

compared with delivery without carrier after i.p. injection. 

25.3.3 POLYPROPYLENIMINE DENDRIMERS 

Polypropylenimine (PPI) dendrimer is also a commercially available cationic dendrimer, but 

much less interest has been taken in its application to gene delivery systems compared to 

PAMAM dendrimer. Since Buhleier et al. synthesized low molecular weight PPI dendrimer in 

1978, 1 synthesis of the dendrimer has been improved through the Michael addition of amines to 

acrylonitrile, followed by the reduction of the nitriles to primary amines. 82 Diaminobutane (DAB) 

and diaminoethane cores are most commonly used for PPI dendrimer synthesis. 

The potential of PPI dendrimers as gene delivery carriers was first compared with other cationic 

polymers including branched PEI (25 kDa), linear PEI (22 kDa), and PAMAM dendrimer and 

others in Cos-7 cell line. 83 Although PPI dendrimer G5 (Astramole-64) showed very low transfection 

efficiency, MTT assay demonstrated that its cytotoxicity was relatively low compared with 

other polymers in treating polyplexes or free polycations. However, PPI dendrimers (generations 1– 

5) with DAB core were re-evaluated as efficient gene delivery systems by Uchegbu and coworkers. 

84 They reported that DNA binding increased with dendrimer generation through molecular 

modeling, and experimental data and cytotoxicity followed the trend DAB 64 (terminal amine 

number)ODAB 32ODAB 16ODOTAPODAB 4ODAB 8, whereas transfection efficacy 

followed the trend DAB 8ZDOTAPZDAB 16ODAB 4ODAB 32ZDAB 64. They concluded 

that the generation 2 polypropylenimine dendrimer (DAB 8) with a low level of cytoxicity gave 

optimum gene transfer activity. 

PPI dendrimers were also chemically modified for improved physiological properties. Park 

and coworkers investigated ternary complex of PPI dendrimer linked with DABs to its periphery 

(PPI–DAB), DNA, and cucurbituril (CB) for a non-covalent strategy in developing a gene 

delivery carrier. 85 CB is a large-cage compound composed of glycoluril units interconnected 

with methylene bridges, and it is able to form stable pseudorotaxanes with strings derived 

from DAB moeities of PPI–DAB through multiple non-covalent interactions. 86,87 The DNA 

binding capacities of PPI, PPI–DAB, and PPI–DAB/CB decreased with increasing generation 

and increased in the order, PPIOPPI–DABRPPI–DAB/CB. Fluorescence experiment using 

DAB-tethered acridine indicated that there was no change in the CB beads threaded on 

PPI–DAB after the addition of DNA to the PPI–DAB/CB binary complex. Importantly, 

in vitro transfection results showed that CB had no negative influence on the transfection efficiency 

of polymer/DNA binary complex, and the ternary complex was a highly active form of 

gene delivery vector, suggesting that this promising system can be applied to other non-covalent 

molecular recognitions, and it would be possible to create a really functional gene delivery carrier 

if the functionalized-CB were synthesized. Recently, another approach to developing gene 

delivery carriers with PPI dendrimers involved methyl quaternary ammonium derivatives of 

PPI dendrimers with varying generation. 88 Quaternization of DAB-cored PPI dendrimer (DAB 

8) with eight surface amines (Q 8) proved to be critical in enhancing DNA binding and resulted in 

improved colloidal stability and vector tolerability on i.v. injection of the polyplex. 

Quaternization was also found to improve in vitro cell biocompatiblity of DAB 16 and DAB 32 

polyplexes. Interestingly, the intravenous administration of DAB 16 and Q8 polyplexes resulted in 

liver-targeted gene expression as opposed to the lung-targeted gene expression obtained with the 


502 

control polymer, Exgen 500. This intrinsic targeting ability without the need for targeting ligands 

may be exploited for the targeted delivery of genes in the treatment of liver diseases. Also, Dufes 

et al. synthesized a vector system based on PPI dendrimer and investigated its transfection in tumors 

after i.v. injection. 89 The systemic tumor necrosis factor alpha (TNFalpha) gene delivery by the 

dendrimer was found to be efficient and non-toxic in the treatment of various established carcinomas. 

Interestingly, these dendrimers and other common polymeric transfection agents also 

exhibited plasmid-independent anti-tumor activity, ranging from pronounced growth retardation 

to complete tumor regression, indicating that the combination of the pharmacological active 

dendrimer and gene would be more potent. 

25.3.4 OTHER DENDRIMERS 


Dendritic poly(L-lysine)s (DPKs) of several generations were synthesized and investigated as gene 

transfection reagents. 90 The DPKs of the fifth and sixth generation showed efficient and serumtolerant 

gene transfection ability into several cell lines without significant cytotoxity. Several 

uptake inhibitor experiments suggested that the uptake of the DNA complex of DPK is mediated 

by the endocytosis pathway, especially the macropinocytotic process. In addition, terminal amino 

acids of DPK G6 were replaced by arginines and histidines to investigate the effect of substituting 

terminal cationic groups on the gene delivery. 91 Arginine-conjugated polylysine dendrimer 

(KGR6) showed higher transfection efficiency than unmodified KG6. In contrast, histidine-conjugated 

polylysine dendrimer (KGH6) showed high transfection efficiency only after polyplex 

formation under acidic conditions because of imdazole groups with a low pKa value of histidine 

residues. pH-dependent complex formation and transfection of KGH6 showed the potential for a 

functional gene delivery system. 

Choi et al. synthesized PEG–dendrimer conjugates PEG–poly(L-lysine) dendrimers and poly 

(L-lysine) dendrimer–PEG–poly(L-lysine) dendrimers (PLLD–PEG–PLLD) and characterized 

their self-assembly with plasmid DNA. 92,93 PEG was used as a core and poly(L-lysine) dendrimers 

were extended via a divergent peptide synthesis method. The dendrimers could form nanosized 

and water-soluble complexes with plasmid DNA in a generation-dependent manner and showed 

low cytotoxicity, but their transfection was not investigated. Thereafter, Kim et al. prepared 

triblock copolymers and PAMAM–PEG–PAMAM dendrimers for gene delivery (Figure 25.7). 94 

The copolymers showed the generation-dependent formation of nanosized complexes with 

plasmid DNA as PLLD–PEG–PLLD and high water-solubility because of their PEG cores. The 

G5 copolymer was found to possess low cytotoxicity and a much enhanced transfection ability in 

comparison to PLLD–PEG–PLLD in HepG2 cells and 293 cells. 

Joester et al. developed amphiphilic dendrimers with rigid tolane cores for gene delivery 

(Figure 25.8). 95 The dendrimers showed hydrophobicity, nominal charge-dependent DNA 

binding affinity, low cytotocity, and very high transfection efficiency superior to commercial 

transfection agents. However, the significant reduction of transfection efficiency with serum is 

thought to need research into the interaction of the dendrimer with anionic serum proteins. 

A series of water-soluble dendrimers with a phosphoramidothioate backbone (P-dendrimers) 

were prepared for gene delivery. 96 The polymers showed rather moderate cytotoxicity toward 

HeLa, HEK 293, and HUVEC cells, and they efficiently delivered oligodeoxyribonucleotide into 

HeLa cells in serum-containing medium. 

Hussain et al. developed another gene delivery system based on anionic dendrimer by directly 

conjugating ODNs to pentaerythritol-based phosphoroamidite dendrimer. 97 The cellular uptake 

of the ODN–dendrimer conjugates was up to 4-fold greater than for naked ODN in cancer cells. 

The ODN–dendrimer conjugates were found to effectively inhibit cancer cell growth, showing a 

marked knockdown in EGFR protein expression. 

Recently, Nishiyama et al. reported the application of photochemical internalization (PCI)mediated 

gene delivery in vivo using dendrimeric photosentisizer. 98 PCI is that the cytoplasmic 



H2N NH2 NH H 

2 

2N NH 

H NH 

2N 

NH 

2 

NH 

NH NH 

2 

2 

NH 

NH 

H NH 

H NH 

2N N 

2 

N 

NH 

H NH 

2N HN 

NH 

2 

H N 

HN 

NH 

NH 

N 

H N 

N 

N 

H HN 

NH 

NH 

2N 

2 

N 

H NH 

2N 2 

HN 

NH 

N 

H N 

N 

HN N 

NH 

H 

H NH 

2N N 

NH 

2 

HN 

NH 

HN 

NH 

N 

N 

H NH2 2N N 

H 

NH 

NH 

HN 

HN 

NH 

N 

NH NH 

2 N 

2 

NH 

H NH2 2N HN 

NH 

NH 

NH 

NH H 

2 

2N H NH NH 

2N H2N 2 

2 

HN 

H2N NH2 HN 

NH HN 

HN O O 

O O 

O 

O HN 

N 

N 

O 

O 

N 

N 

O 

O 

HN 

HN 

O 

O 

O 

O 

N 

N 

N 

O 

O 

O 

HN 

O 

O 

O 

O 

N 

O 

HN 

N 

O O 

N O O N H 

O 

N 

O 

O 

O 

O 

O 

O 

O 

O N 

O 

O 

O 

O 

O 

N 

O O 

N O O N H 

N 

HN 

O 

N H O 

O 

O 

O 

O 

O 

N 

O 

O 

O 

N 

O 

O 

N 

HN 

H 

O 

O 

N 

O 

O 

N 

O HN 

O 

O O 

O 

O 

HN 

HN 

H 

H 

N 

FIGURE 25.7 Molecular structure of PAMAM–PEG–PAMAM G4. 

delivery of macromolecular compounds is enhanced by the photochemical disruption of the endosomal 

membrane in using light and a hydrophilic photosensitizer. 99 They prepared ternary 

complexes having cores of pDNA/cationic peptide polyplexes enveloped with anionic dendrimer 

phthalocyanine (DPc). The complexes showed greatly enhanced transfection efficiency and low 

H 

N 

O 

O 

HN 

NH 

FIGURE 25.8 Molecular structure of amphiphilic dendrimer with tolane core. 


O 

N 

H 

O 

O 

N 

H 

O 

+ H3 N NH3 + 

HN 

O 

H 

N 

O 

NH 

NH 3 + 

NH 3 + 

NH 3 + 

NH 3 + 

NH 3 + 

NH 3 + 

NH 3 +

504 

photocytotoxicity in HeLa cells after photoirradiation, and they achieved significant gene 

expression in conjunctival tissue of rat eyes in contrast to ExGen500 that is one of the most 

efficient vectors. 

25.4 CONCLUSION 

In the past decade, various dendrimers have been developed and investigated for drug and gene 

delivery systems because of their unique structural properties such as multivalent and controllable 

functionalities on the periphery and highly branched and globular architectures as shown in this 

review. It is apparent that the application of dendrimers to drug and gene delivery systems has 

achieved considerably successful results, and much of the work reported in medical and pharmaceutical 

journals has demonstrated the growing interest in dendrimers. 

Although it is not presented in this chapter, the application of dendrimers has been extended in 

other medical fields such as diagnostics, antiviral, or antibacterial therapy. PAMAM dendrimerbased 

Gd III chelates were synthesized for use as a MRI contrast agent. 100,101 MRI is a powerful 

technique in medical diagnostics and used to visualize organs and blood vessels. The dendrimer 

chelates showed excellent MRI images and long blood circulation times in in vivo experiments. 

Also, chemically modified polylysine dendrimers and other dendrimers were synthesized for useful 

antiviral or antibacterial drugs. 102–105 In addition, numerous dendrimeric glycans or glycosylated 

dendrimers were shown to efficiently interact with natural carbohydrates receptors and were 

focused in their potential as microbial anti-adhesins, microbial toxin antagonists, and anti-cancer 

drugs. 106,107 

However, research about stable and safe biodistribution and efficient tissue localization as well 

as the biological function and fate of dendrimers is not well-established. Therefore, it is thought to 

be very important to explore the physiological safety and pharmacokinetics of developed dendrimers 

over a vaster range. It is also expected that a more convenient methodology to synthesizing 

novel dendrimers giving high water-solubility, good biocompatibility, and perfect structure will be 

designed for efficient application of dendrimers to biological systems. 


This work was supported by the National R&D Program Grant of The Ministry of Commerce, 

Industry and Energy (M10422010006-04N2201-00610), and the Gene Therapy Project of The 

Ministry of Science and Technology (M1053403004-05N3403-00410). The authors also wish to 

acknowledge helpful assistance received from Jung-un Baek. 

REFERENCES 


1. Buhleier, E., Wehner, W., and Vögtle, F., Cascade- and nonskid-chain-like” syntheses of molecular 

cavity topologies, Synthesis, 155–158, 1978. 

2. Zeng, F. and Zimmerman, S. C., Dendrimers in supramolecular chemistry: From molecular recognition 

to self-assembly, Chem. Rev., 97, 1681–1712, 1977. 

3. Fischer, M. and Vögtle, F., Dendrimers: From design to application—a progress report, Angew. 

Chem. Int. Ed., 38, 884–905, 1999. 

4. Stiriba, S. E., Frey, H., and Haag, R., Dendritic polymers in biomedical applications: From potential 

to clinical use in diagnostics and therapy, Angew. Chem. Int. Ed., 41, 1329–1334, 2002. 

5. Patri, A. K., Majoros, I. J., and Baker, J. R. Jr., Dendritic polymer macromolecular carriers for drug 

delivery, Curr. Opin. Chem. Biol., 6, 466–471, 2002. 

6. Gilles, E. R. and Frechet, J. M. J., Dendrimers and dendritic polymers in drug delivery, Drug Discov. 

Today, 10, 35–43, 2005. 



7. Blume, G. and Cevc, G., Liposomes for sustained drug release in vivo, Biochim. Biophys. Acta, 1029, 

97, 1990. 

8. Lian, T. and Ho, R. J. Y., Trends and developments in liposome drug delivery systems, J. Pharm. 

Sci., 90, 680, 2001. 

9. Rosler, A., Vandermeulen, G. W. M., and Klok, H. A., Advanced drug delivery devices via selfassembly 

of amphiphilic block copolymers, Adv. Drug Deliv. Rev., 53, 95–108, 2001. 

10. Newkome, G. R. et al., Unimolecular micelles, Angew. Chem. Int. Ed. Engl., 30, 1178–1180, 1991. 

11. Jansen, J. F. G. A., de Brabander-van den Berg, E. M. M., and Meijer, E. W., Encapsulation of guest 

molecules into a dendritic box, Science, 266, 1226–1229, 1994. 

12. Jansen, J. F. G. A., Meijer, E. W., and de Brabander-van den Berg, E. M. M., The dendritic box: 

Shape-selective liberation of encapsulated guests, J. Am. Chem. Soc., 117, 4417–4418, 1995. 

13. Baars, M. W. P. L. et al., The localization of guests in water-soluble oligoethyleneoxy-modified 

poly(propylene imine) dendrimers, Angew. Chem. Int. Ed., 39, 1285–1288, 2000. 

14. Sideratou, Z., Tsiourvas, D., and Paleos, C. M., Solubilization and release properties of PEGylated 

diaminobutane poly(propylene imine) dendrimers, J. Colloid Interface Sci., 242, 272–276, 2001. 

15. Sideratou, A., Tsiourvas, D., and Paleos, C. M., Quaternized poly(propylene imine) dendrimers as 

novel pH-sensitive controlled-release systems, Langmuir, 16, 1766–1769, 2000. 

16. Tomalia, D. A. et al., A new class of polymers: Starbust-dendritic macromolecules, Polym. J., 17, 

117–132, 1985. 

17. Twyman, L. J. et al., The synthesis of water soluble dendrimers, and their application as possible drug 

delivery systems, Tetrahedron Lett., 40, 1743–1746, 1999. 


their ability to encapsulate anticancer drugs, Bioconjug. Chem., 11, 910–917, 2000. 

19. Bhadra, D. et al., A PEGylated dendritic nanoparticle carrier of fluorouracil, Int. J. Pharm., 257, 

111–124, 2003. 

20. Haba, Y. et al., Synthesis of biocompatible dendrimers with a peripheral network formed by linking 

of polymerizable groups, Polymer, 46, 1813–1820, 2005. 

21. Liu, M., Kono, K., and Frechet, J. M. J., Water-soluble unimolecular micelles: Their potential as 

drug delivery agents, J. Control. Release, 65, 121–131, 2000. 

22. Gillies, E. R., Jonsson, T. B., and Frechet, J. M. J., Stimuli-responsive supramolecular assemblies of 

linear-dendritic copolymers, J. Am. Chem. Soc., 126, 11936–11943, 2004. 

23. Namazi, H. and Adeli, M., Dendrimers of citric acid and poly(ethylene glycol) as the new drugdelivery 

agents, Biomaterials, 26, 1175–1183, 2005. 

24. Roberts, J. C. et al., Using starbust dendrimers as linker molecules to radiolabel antibodies, 

Bioconjug. Chem., 1, 305–308, 1990. 

25. Wu, C. et al., Metal–chelate–dendrimer–antibody constructs for use in radioimmunotherapy and 

imaging, Bioorg. Med. Chem. Lett., 4, 449–454, 1994. 

26. Patri, A. K. et al., Antibody–dendrimer conjugates for targeted prostate cancer therapy, Polym. 

Mater. Sci. Eng., 86, 130, 2002. 

27. Duncan, R. and Malik, N., Dendrimers: Biocompatibility and potential for delivery of anticancer 

agents, Proc. Int. Symp. Control. Release Bioact. Mater., 23, 105–106, 1996. 

28. Malik, N., Evagorou, E. G., and Duncan, R., Dendrimer–platinate: A novel approach to cancer 

chemotherapy, Anticancer Drugs, 10, 767–776, 1999. 

29. Zhuo, R. X., Du, B., and Lu, Z. R., In vitro release of 5-fluorouracil with cyclic core dendritic 

polymer, J. Control. Release, 57, 249–257, 1999. 

30. Wang, D. et al., Synthesis of starlike N-(2-hydroxypropyl)methacrylamide copolymers: Potential 

drug carriers, Biomacromolecules, 1, 313–319, 2000. 

31. Wiwattanapatapee, R., Lomlim, L., and Saramunee, K., Dendrimers conjugates for colonic delivery 

of 5-aminosalicylic acid, J. Control. Release, 88, 1–9, 2003. 

32. Quintana, A. et al., Design and function of a dendrimer-based therapeutic nanodevice targeted to 

tumor cells through the folate receptor, Pharm. Res., 19, 1310–1316, 2002. 

33. Tansey, W. et al., Synthesis and characterization of branched poly(L-glutamic acid) as a biodegradable 

drug carrier, J. Control. Release, 94, 39–51, 2004. 

34. Kono, K., Liu, M., and Fréchet, J. M. J., Design of dendritic macromolecules containing folate or 

methotrexate residues, Bioconjug. Chem., 10, 1115–1121, 1999. 


506 


35. Liu, M., Kono, K., and Fréchet, J. M. J., Water-soluble dendrimer–polyethylene glycol starlike 

conjugates as potential drug carriers, J. Polym. Sci. A, 37, 3492–3503, 1999. 

36. Choe, Y. H. et al., Anticancer drug delivery systems: Multi-loaded N 4 -acyl poly(ethylene glycol) 

prodrugs of ara-C. II. Efficacy in ascites and solid tumors, J. Control. Release, 79, 55–70, 2002. 

37. Schiavon, O. et al., PEG–ara-C conjugates for controlled release, Eur. J. Med. Chem., 39, 123–133, 2004. 

38. Padilla De Jesu’s, O. L. et al., Polyester dendritic systems for drug delivery applications: In vitro and 

in vivo evaluation, Bioconjug. Chem., 13, 453–461, 2002. 

39. de Groot, F. M. H. et al., “Cascade-release dendrimers” liberation all end groups upon a single 

triggering event in the dendritic core, Angew. Chem. Int. Ed., 42, 4490–4494, 2003. 

40. Amir, R. J. et al., Self-immolative dendrimers, Angew. Chem. Int. Ed., 42, 4494–4499, 2003. 

41. Shamis, M., Lode, H. N., and Shabat, D., Bioactivation of self-immolative dendritic prodrugs by 

catalytic antibody 38C2, J. Am. Chem. Soc., 126, 1726–1731, 2004. 

42. Hawthorne, M. F., The role of chemistry in the development of boron neutron capture therapy of 

cancer, Angew. Chem. Int. Ed. Engl., 32, 950–984, 1993. 

43. Barth, R. F. et al., Boronated starbust dendrimer–monoclonal antibody immunoconjugates: Evaluation 

as a potential delivery system for neutron capture therapy, Bioconjug. Chem., 5, 58–66, 1994. 


capture therapy of brain tumors, Bioconjug. Chem., 7, 7–15, 1996. 

45. Yang, W. et al., Intratumoral delivery of boronated epidermal growth factor for neutron capture 

therapy of brain tumors, Cancer Res., 57, 4333–4339, 1997. 

46. Wu, G. et al., Site-specific conjugation of boron-containing dendrimers to anti-EGF receptor monoclonal 

antibody cetuximab(IMC-C225) and its evaluation as a potential delivery agent for neutron 

capture therapy, Bioconjug. Chem., 15, 185–194, 2004. 


dendrimers as potential agents for neutron capture therapy, Bioconjug. Chem., 14, 158–167, 2003. 

48. Qualmann, B. et al., Synthesis of boron-rich lysine dendrimers as protein labels in electron 

microscopy, Angew. Chem. Int. Ed. Engl., 35, 909–911, 1996. 

49. Matsumoto, T., Transport calculations of depth–dose distributions for gadolinium neutron capture 

therapy, Phys. Med. Biol., 37, 155–162, 1992. 

50. Khokhlov, V. F. et al., Neutron capture therapy with gadopentetate dimeglumine: Experiments on 

tumor-bearing rats, Acad. Radiol., 2, 392–398, 1995. 

51. Hofmann, B. et al., Gadolinium neutron capture therapy (GdNCT) of melanoma cells and solid 

tumors with the magnetic resonance imaging contrast agent Gadobutrol, Invest. Radiol., 34, 

126–133, 1999. 

52. Caravan, P. et al., Gadolinium(III) chelates as MRI contrast agents: Structure, dynamics, and applications, 

Chem. Rev., 99, 2293–2352, 1999. 

53. Kobayashi, H. et al., Avidin-dendrimer-(1B4M-Gd)254: A tumor-targeting therapeutic agent for 

gadolinium neutron capture therapy of intraperitoneal disseminated tumor which can be monitored 

by MRI, Bioconjug. Chem., 12, 587–593, 2001. 

54. Dougherty, T. J. et al., Photodynamic therapy, J. Natl. Cancer Inst., 90, 889–905, 1998. 

55. Battah, S. H. et al., Synthesis and biological studies of 5-aminolevulinic acid-containing dendrimers 

for photodynamic therapy, Bioconjug. Chem., 12, 980–988, 2001. 

56. Peng, Q. et al., 5-Aminolevulinic acid-based photodynamic therapy: Principles and experimental 

research, Photochem. Photobiol., 65, 235–251, 1997. 

57. Nishiyama, N. et al., Light-harvesting ionic dendrimer porphyrins as new photosensitizers for photodynamic 

therapy, Bioconjug. Chem., 14, 58–66, 2003. 

58. Zhang, G. D. et al., Polyion complex micelles entrapping cationic dendrimer porphyrin: Effective 

photosensitizer for photodynamic therapy of cancer, J. Control. Release, 93, 141–150, 2003. 

59. Dichtel, W. R. et al., Singlet oxygen generation via two-photon excited FRET, J. Am. Chem. Soc., 

126, 5380–5381, 2004. 

60. Boussif, O. et al., A versatile vector for gene and oligonucleotide transfer into cells in culture and 

in vivo: Polyethylenimine, Proc. Natl Acad. Sci., 92, 7297–7301, 1995. 

61. Rejman, J., Bragonzi, A., and Conese, M., Role of clathrin- and caveolae-mediated endocytosis in 

gene transfer mediated by lipo- and polyplexes, Mol. Ther., 12, 468–474, 2005. 



62. Haensler, J. and Szoka, F. C. Jr., Polyamidoamine cascade polymers mediate efficient transfection of 

cells in culture, Bioconjug. Chem., 4, 372–379, 1993. 

63. Kukowska-Latallo, J. F. et al., Efficient transfer of genetic material into mammalian cells using 

starburst polyamidoamine dendrimers, Proc. Natl Acad. Sci. USA, 93, 4897–4902, 1996. 

64. Harada, Y. et al., Highly efficient suicide gene expression in hepatocellular carcinoma cells by 

Epstein–Barr virus-based plasmid vectors combined with polyamidoamine dendrimer, Cancer 

Gene Ther., 7, 27–36, 2000. 

65. Tanaka, S. et al., Targeted killing of carcinoembryonic antigen (CEA)-producing cholangiocarcinoma 

cells by polyamidoamine dendrimer-mediated transfer of an Epstein–Barr virus (EBV)-based 

plasmid vector carrying the CEA promoter, Cancer Gene Ther., 7, 1241–1250, 2000. 

66. Tang, M. X., Redemann, C. T., and Szoka, F. C. Jr., In vitro gene delivery by degraded polyamidoamine 

dendrimers, Bioconjug. Chem., 7, 703–714, 1996. 

67. Bielinska, A. et al., Regulation of in vitro gene expression using antisense oligonucleotides or 

antisense expression plasmids transfected using starburst PAMAM dendrimers, Nucleic Acids 

Res., 24, 2176–2182, 1996. 

68. Delong, R. et al., J. Pharm. Sci., 86, 762–764, 1997. 

69. Yoo, H. and Juliano, R. L., Enhanced delivery of antisense oligonucleotides with fluorophoreconjugated 

PAMAM dendrimers, Nucleic Acids Res., 28, 4225–4231, 2000. 

70. Maksimenko, A. V. et al., Optimisation of dendrimer-mediated gene transfer by anionic oligomers, 

J. Gene Med., 5, 61–71, 2003. 

71. Roessler, B. J. et al., Substituted b-cyclodextrins interact with PAMAM dendrimer–DNA complexes 

and modify transfection efficiency, Biochem. Biophys. Res. Commun., 283, 124–129, 2001. 

72. Kihara, F. et al., Effects of structure of polyamidoamine dendrimer on gene transfer efficiency of the 

dendrimer conjugate with a-cyclodextrin, Bioconjug. Chem., 13, 1211–1214, 2002. 

73. Wada, K. et al., Improvement of gene delivery mediated by mannosylated dendrimer/a-cyclodextrin 

conjugates, J. Control. Release, 104, 397–413, 2005. 

74. Choi, J. S. et al., Enhanced transfection efficiency of PAMAM dendrimer by surface modification 

with L-arginine, J. Control. Release, 99, 445–456, 2004. 

75. Tung, C. H. and Weissleder, R., Arginine containing peptides as delivery vectors, Adv. Drug Deliv. 

Rev., 55, 281–294, 2003. 

76. Kang, H. et al., Tat-conjugated PAMAM dendrimers as delivery agents for antisense and siRNA 

oligonucleotides, Pharm. Res., 22, 2099–2106, 2005. 

77. Lee, J. H. et al., Polyplexes assembled with internally quaternized PAMAM-OH dendrimer and 

plasmid DNA have a neutral surface and gene delivery potency, Bioconjug. Chem., 14, 

1214–1221, 2003. 

78. Hudde, T. et al., Activated polyamidoamine dendrimers, a non-viral vector for gene transfer to the 

corneal endothelium, Gene Ther., 6, 939–943, 1999. 

79. Wang, Y. et al., DNA/dendrimer complexes mediate gene transfer into murine cardiac transplants ex 

vivo, Mol. Ther., 2, 602–608, 2000. 

80. Kukowska-Latallo, J. F. et al., Intravascular and endobronchial DNA delivery to murine lung tissue 

using a novel, nonviral vector, Hum. Gene Ther., 11, 1385–1395, 2000. 

81. Sato, N. et al., Tumor targeting and imaging of intraperitoneal tumors by use of antisense oligo-DNA 

complexed with dendrimers and/or avidin in mice, Clin. Cancer Res., 7, 3606–3612, 2001. 

82. de Brabander-van den Berg, E. M. M. and Meijer, E. W., Poly(propylene imine) dendrimers: Largescale 

synthesis by hetereogeneously catalyzed hydrogenations, Angew. Chem. Int. Ed. Engl., 32, 

1308–1311, 1993. 

83. Gebhart, C. L. and Kabanov, A. V., Evaluation of polyplexes as gene transfer agents, J. Control. 

Release, 73, 401–416, 2001. 

84. Zinselmeyer, B. H. et al., The lower-generation polypropylenimine dendrimers are effective genetransfer 

agents, Pharm. Res., 19, 960–967, 2002. 

85. Lim, Y. b. et al., Self-assembled ternary complex of cationic dendrimer, cucurbituril, and 

DNA: Noncovalent strategy in developing a gene delivery carrier, Bioconjug. Chem., 13, 

1181–1185, 2002. 


508 


86. Whang, D. et al., Self-assembly of a polyrotaxane containing a cyclic “bead” in every structural unit 

in the solid state: Cucurbituril molecules threaded on a one-dimensional coordination polymer, 

J. Am. Chem. Soc., 118, 11333–11334, 1996. 

87. Lee, J. W. et al., Novel pseudorotaxane-terminated dendrimers: Supramolecular modification of 

dendrimer periphery, Angew. Chem. Int. Ed., 40, 746–749, 2001. 

88. Schatzlein, A. G. et al., Preferential liver gene expression with polypropylenimine dendrimers, 

J. Control. Release, 101, 247–258, 2005. 

89. Dufes, C. et al., Synthetic anticancer gene medicine exploits intrinsic antitumor activity of cationic 

vector tocure established tumors, Cancer Res., 65, 8079–8084, 2005. 

90. Ohsaki, M. et al., In vitro gene transfection using dendritic poly(L-lysine), Bioconjug. Chem., 13, 

510–517, 2002. 

91. Okuda, T. et al., Characters of dendritic poly(L-lysine) analogues with the terminal lysines replaced 

with arginines and histidines as gene carriers in vitro, Biomaterials, 25, 537–544, 2004. 

92. Choi, J. S. et al., Poly(ethylene glycol)-block-poly(L-lysine) dendrimer: Novel linear polymer/dendrimer 

block copolymer forming a spherical water-soluble polyionic complex with DNA, Bioconjug. 

Chem., 10, 62–65, 1999. 

93. Choi, J. S. et al., Synthesis of a barbell-like triblock copolymer, poly(L-lysine) dendrimer-blockpoly(ethylene 

glycol)-block-poly(L-lysine) dendrimer, and its self-assembly with plasmid DNA, 

J. Am. Chem. Soc., 122, 474–480, 2000. 

94. Kim, T. I. et al., PAMAM–PEG–PAMAM: Novel triblock copolymer as a biocompatible and 

efficient gene delivery carrier, Biomacromolecules, 5, 2487–2492, 2004. 

95. Joester, D. et al., Amphiphilic dendrimers: Novel self-assembling vectors for efficient gene delivery, 

Angew. Chem. Int. Ed., 42, 1486–1490, 2003. 

96. Maszewska, M. et al., Water-soluble polycationic dendrimers with a phosphoramidothioate backbone: 

Preliminary studies of cytotoxicity and oligonucleotide/plasmid delivery in human cell 

culture, Oligonuleotides, 13, 193–205, 2003. 

97. Hussain, M. et al., A novel anionic dendrimer for improved cellular delivery of antisense oligonucleotides, 

J. Control. Release, 99, 139–155, 2004. 

98. Nishiyama, N. et al., Light-induced gene transfer from packaged DNA enveloped in a dendrimeric 

photosensitizer, Nat. Mater., 4, 934–941, 2005. 

99. Høgest, A. et al., Photochemical internalization in drug and gene delivery, Adv. Drug Deliv. Rev., 56, 

95–115, 2004. 

100. Wiener, E. C. et al., Dendrimer-based metal chelates: A new class of magnetic resonance imaging 

contrast agents, Magn. Reson. Med., 31, 1–8, 1994. 

101. Bryant, L. H. Jr. et al., Synthesis and relaxometry of high-generation (GZ5, 7, 9, and 10) PAMAM 

dendrimer–DOTA–gadolinium chelates, J. Magn. Reson. Imaging, 9, 348–352, 1999. 

102. Bourne, N. et al., Dendrimers, a new class of candidate topical microbicides with activity against 

herpes simplex virus infection, Antimicrob. Agents Chemother., 44, 2471–2474, 2000. 

103. Witvrouw, M. et al., Potent anti-HIV (type 1 and type 2) activity of polyoxometalates: Structure– 

activity relationship and mechanism of action, J. Med. Chem., 43, 778–783, 2000. 

104. Landers, J. J. et al., Prevention of influenza pneumonitis by sialic acid-conjugated dendritic polymers, 

J. Infect. Dis., 186, 1222–1230, 2002. 

105. Chen, C. Z. S. and Cooper, S. L., Recent advances in antimicrobial dendrimers, Adv. Mater., 12, 

843–846, 2000. 

106. Veprek, P. and Jezek, J., Peptide and glycopeptide dendrimers. Part II, J. Pept. Sci., 5, 203–220, 

1999. 

107. Andre, S. et al., Wedgelike glycodendrimers as inhibitors of binding of mammalian galectins to 

glycoproteins, lactose maxiclusters, and cell surface glycoconjugates, Chembiochem, 2, 822–830, 

2001. 


26 

CONTENTS 

Dendritic Nanostructures 


Ashootosh V. Ambade, Elamprakash N. Savariar, 

and S. (Thai) Thayumanavan 

26.1 Introduction ....................................................................................................................... 509 

26.2 Dendrimer–Drug Conjugates ............................................................................................ 511 

26.3 Encapsulation of the Drug in a Dendrimer....................................................................... 514 

26.4 Polymeric Drug Carriers Based on Dendrimers ............................................................... 517 

26.5 Summary............................................................................................................................ 518 

References..................................................................................................................................... 518 


Nanomaterials have emerged from academic curiosity and are being applied in many diverse areas 

such as information technology and biotechnology. 1–4 Synthetic nanomaterials based on macromolecules 

are relevant in biomedical applications because their size scale is comparable with that of 

biomacromolecules. 5 Dendrimers provide a prominent example of such synthetic macromolecules 

that have attracted great attention in recent years because of their potential in drug delivery among 

other promising applications. 6–12 Dendrimers are highly branched polymers of a few nanometer 

dimensions that possess a globular shape at high molecular weights. 13,14 These are synthesized by 

repetitive organic transformations and, as a result, are almost monodisperse. An interesting structural 

feature of dendrimers is the spatially isolated core in the dendritic interior that is reminiscent 

of site-isolation in biomolecules such as proteins. The branches emanating from this core provide 

cavities in the interior that can be exploited for guest encapsulation. Every layer of branches that is 

added to the core during the synthesis is termed as a generation. The highly branched nature of 

dendrimers leads to large number of peripheral groups. These surface functionalities can be 

modified by attachment of drugs or targeting ligands to meet specific applications that often 

involve controlled interaction with cell walls and other biological targets. A representative structure 

is shown in Figure 26.1 that illustrates the typical features—core, intermediate layers, and peripheral 

groups—of a dendrimer. The dendritic structure in terms of type of repeat units, size of interior 

cavities, and type of peripheral groups can be tuned by using either convergent or divergent 

synthetic strategy. 15,16 It is then obvious that such versatile molecules are being intensely 

pursued in cancer therapeutics. 17,18 

A major problem with anti-cancer agents is that they indiscriminately attack healthy cells as 

well as cancerous cells. These harmful side effects can be reduced by developing a drug delivery 

vehicle that is specific to tumor cells. This can be achieved by employing a strategy called active 


509

510 

H 2N 

H 2N 

H 

N 

O 

H 

N 

Intermediate layers 

(dendritic interior) 

H 2N 

N 

H 

H 2N 

H 2N 

O 

O 

N 

HN 

N 

O 

O 

NH 

H 

N 

O 

NH 

H 2N 

O 

N 

O 

HN 

N 

N 

NH 2 

NH 

O 

NH 

O 

NH 

N 

O 

NH O 

HN O 

O 

HN 

NH 2 

N 

Core 

targeting wherein functionalities that respond to over-expressed receptors on tumor cells are 

attached to the drug carrier. Dendrimers provide an opportunity in this direction in the form of a 

large number of tunable peripheral groups where a variety of functionalities can be incorporated 

(Figure 26.2). Another strategy is the passive targeting that relies on the well-known enhanced 

permeation and retention (EPR) effect. This effect, first described by Maeda et al., 19 takes advantage 

of poorly formed (leaky) vasculature of solid tumors that allows accumulation of polymer–drug 

HN 

O 

N 

O 

HN 

O 

HN 

O 

NH 

O 

N 

N 

N 

O 

NH 

H 2N 

Peripheral amine groups 

N 

NH 

NH 2 

HN 

H 

N 

O 

O 

NH 

H 2N 

O 

O 

NH 

N 

O 

O 

N 

NH 2 

H 

N 

H 

N 

O 

O 

N 

H 

NH 

FIGURE 26.1 Structural features of a dendrimer are illustrated with a structure of amine-terminated PAMAM 

dendrimer. 

= Drug molecule 

= Targetingmoiety 

= Functionality for imaging 

FIGURE 26.2 A dendrimer with multiple functionalities attached to it is illustrated by a cartoon. 



NH 2 

NH 2 

NH 2 

NH 2

Dendritic Nanostructures for Cancer Therapy 511 

conjugates ranging in size between 10 and 500 nm within tumors than of free drugs. The 

polymeric molecules are retained following the accumulation due to their larger size whereas 

free drug molecules are easily eliminated from the cells. Dendrimers have an advantage for this 

strategy as well because drugs can be physically encapsulated into the interior of nanometer-sized 

dendrimers, thereby increasing the uptake of drugs into tumors. Before describing the various ways 

in which dendritic structure has been utilized for drug delivery in cancer therapy, it is useful to 

discuss the techniques that are employed for this purpose. 

Photodynamic therapy (PDT), also known as photoradiation therapy or photochemotherapy, 

involves the use of a light-activated drug called photosensitizer that can be administered orally or 

intraperitoneally, or it can be applied to the skin. 20,21 When light of a specific wavelength that is 

absorbed by the dye is shined on the targeted tissue, the photosensitizer is excited from ground 

singlet state to excited singlet state. It then undergoes intersystem crossing to a longer-lived excited 

triplet state that reacts with the ground triplet state of molecular oxygen present in the tissues. This 

energy transfer creates highly-reactive excited singlet-state oxygen species. The singlet oxygen 

reacts destructively with nearby biomolecules and induces cell death either by apoptosis or 

necrosis. PDT, however, has limitations because it is applicable only in areas where light can reach. 

Neutron capture therapy (NCT) is another promising technique for treatment of cancer. 22 It is 

based on the nuclear reaction that occurs when a stable isotope, for example, boron 10 B, is irradiated 

with low energy or thermal neutrons to produce high-energy alpha particles ( 4 He nuclei) and 7 Li 

ions. The combined path length of these two particles is about 12 mm that is close to the diameter of 

a cell. Therefore, if the tumor cells can be supplied with a high concentration of 10 B atoms, the 

effect of the radiation will be limited to tumor cells, and the normal cells may be spared. This is 

particularly important in cancer therapy where prevention of indiscriminate destruction of healthy 

cells is a challenge. 157 Gd is another isotope of interest because of its large neutron capture cross 

section. Success of the above two therapies depends on localizing sufficient concentration of the 

photosensitizer or radioactive material in the tumor cells while having negligible accumulation in 

the healthy cells. 

Magnetic resonance imaging (MRI) is one of the most widely used imaging techniques that has 

greatly improved medical diagnosis. 23 It is based on the relaxation properties of hydrogen nuclei in 

water in body tissues when placed in a uniform magnetic field. In clinical practice, for detection of 

cancer in particular, it uses the fact that the relaxation times of protons in cancerous and normal 

tissues markedly differ. One of the advantages of MRI over other imaging techniques is that it is a 

non-radiative technique and harmless to the patient. 

This chapter will focus on the structural aspects of dendritic molecules that have been used to 

design efficient and multifunctional drug delivery vehicles in cancer therapy with the help of 

selected illustrative examples. There are only few cases where these macromolecules have been 

studied in vivo, i.e., in animal models, and the biological studies will be discussed only cursorily. 

26.2 DENDRIMER–DRUG CONJUGATES 

The term dendrimer–drug conjugate is mainly applied when a drug molecule is covalently linked to 

a dendrimer. The drug can be covalently attached to a dendrimer on the periphery or at the core and 

quite rarely at the branching points, i.e., in the interior layers. Attachment of a drug to multiple 

peripheral groups in a dendrimer greatly increases the effective concentration of a drug once it is 

released at the targeted site. This is particularly advantageous for the use of prodrugs. In the case of 

PDT, however, it is desirable to isolate the drug molecule to prevent fluorescence quenching 

because of aggregation, and this can be achieved by attaching the photosensitizer to the core of 

a dendrimer. A dendrimer–drug conjugate can provide certain advantages over conventional polymeric 

drug delivery vehicles because dendrimers are monodisperse, structurally defined 

macromolecules with a controlled molecular weight and size. 


512 

O 2N 

O 

HN 

O 

O 

O 

O 

OR 

O 

HN O 

O 

O 

O 

OR 

OR 

O 

OR 

O 

R = Ph 

O OH 

When a drug is attached to peripheral groups of a dendrimer, the linker between drug and 

dendrimer is of crucial importance because the drugs need to be efficiently released in active form at 

the targeted site. Advantage is taken of the acidic or reducing environment near the tumor cells by 

attaching the drugs to dendritic periphery via acid-labile or disulphide (susceptible to reduction) 

linker (see Section 26.4 for examples). In this case, each drug is separately cleaved over a long 

period of time. An elegant design was reported in this direction wherein a single triggering event at 

the core results in disassembly of the dendrimer and release of all the paclitaxel molecules attached 

to the periphery. 24 The drug release in these dendrimers incorporating a masked 4-aminocinnamyl 

alcohol linker is activated by reduction of nitro to amino group under mild reducing conditions 


Poly(amidoamine) (PAMAM) dendrimers have been widely used for preparation of 

dendrimer–drug conjugates because of easily modifiable terminal amine groups, commercial availability 

of higher generation dendrimers, and biocompatibility. 25–27 For example, 

methylprednisolone (MP) was attached to PAMAM periphery through a glutaric acid linker. 28 

Because the MP–glutaric acid compound was attached to a preformed dendrimer, only 12 drug 

molecules (32 wt%) could be conjugated to a dendrimer that has 32 terminal moieties as a result of 

steric hindrance. However, considering that MP is a bulky steroidal drug, this was a significant 

increase over earlier payloads. 

Monoclonal antibodies (mAb) are useful for targeted delivery because these can specifically 

bind to certain receptors over-expressed on tumor cells. Antibody–PAMAM conjugates have been 

prepared by attaching J591 anti-PSMA antibody to a fluorescein-labeled dendrimer for targeted 

delivery in prostate cancer therapy. It was based on the fact that prostate-specific membrane antigen 

(PSMA) is over-expressed on tumor cells than on normal tissue. 29 Attachment of multiple functionalities 

for targeting (folic acid) and imaging (fluorescein) along with an anti-cancer drug 

(methotrexate) to a dendrimer periphery has also been demonstrated with PAMAM dendrimers 30 

and is elaborated below with examples on MRI and NCT. Similarly, different PAMAM dendrimers 

attached with fluorescein and folic acid were covalently attached to oligonucleotides and assembled 

via hybridization to afford multifunctional dendrimer clusters as illustrated by a cartoon in 

Figure 26.4. 31 On a different note, carbohydrate moieties have been attached to PAMAM dendrimers 

to study multivalent protein–carbohydrate interactions. This is particularly relevant to cancer 

metastasis that involves these interactions in cellular recognition processes. 32–34 

O 

N 

H 


Ph 

O 

O 

AcO 

H 

OH AcO 

OBz 

Paclitaxel 

Point of attachment to the dendrimer 

FIGURE 26.3 Paclitaxel–dendrimer conjugate with a cascade-release mechanism. The NO 2 group at the focal 

point is converted to NH2 under reducing conditions triggering the cascade-release. 


O


FITC 

PAMAM dendrimer 

FITC 

11 nm 

20 nm 

FITC-functionality for imaging 

FA-functionality for targeting 

PAMAM dendrimers have also been used for preparation of MRI contrast agents. 35 Gd(III) 

complexes were attached to G6-PAMAM dendrimers along with avidin to afford a targeted delivery 

vehicle for NCT. 36 157 Gd provides two advantages. First, it has a large neutron capture cross section 

that can be exploited in NCT, and second, Gd(III) complexes can be used as MRI contrast agents. In 

fact, dendrimers bearing Gd(III) complexes on their periphery have undergone clinical trials as 

MRI contrast agents. 37,38 Therefore, the Gd-dendrimer conjugate described above serves a dual 

purpose, therapeutic as well as imaging. The Gd-dendrimer complex was also used as a contrast 

agent for dynamic MRI to evaluate the perfusion of an antibody viz. herceptin into tumor tissue 

because it has similar size (w10 nm). 39 Similarly, folic acid-conjugated 40 and epidermal growth 

factor (EGF)-conjugated 41 boronated PAMAM dendrimers have been prepared as targeting agents 

in boron NCT. Slow clearance of macromolecular MRI agents is an issue that limits their scope, 

particularly in the investigation of tumor angiogenesis. To address this problem, the avidin chase 

system was employed where biotinylated dendrimeric Gd-complexes were rapidly cleared 

(w2 min) from blood pool to the liver on injection of avidin. 42 Dendrimeric Gd(III) chelates 

with high relaxivity and fast water exchange have been developed to improve upon the currently 

available gadolinium complexes that have slow water exchange rates. 43 In rare examples of interior 

functionalized dendrimers, boron clusters have been incorporated in the intermediate layers of 

water-soluble dendrimers for application in NCT. 44,45 

Porphyrins and phthalocyanines are two important macrocycles for use as photosensitizers in 

PDT because of large p-conjugation domains and high singlet oxygen quantum yields. 21 However, 

large p-conjugation and hydrophobicity in these molecules also lead to a tendency to aggregate that 

reduces their solubility in water and precludes use in biological systems. The aggregation also 

causes self-quenching and decreases the efficiency of singlet oxygen formation. Incorporation of a 

porphyrin or phthalocyanine chromophore at the core of a benzyl ether dendrimer has been shown 

to provide the site isolation and to necessarily prevent the aggregation and self-quenching. Hydrophilic 

groups such as oligoethylene glycol were attached to the periphery of these dendrimers, 

FA 

Oligonucleotide 

FIGURE 26.4 PAMAM clusters with multiple functionalities prepared via DNA hybridization. 


FA

514 

making them useful in biological applications. 46 Oligoethylene glycols are important in biological 

applications because they provide solubility in water, reduce hepatic uptake, and enhance biodistribution. 

It is to be noted that benzyl ether dendrimers are another class of dendrimers, apart from 

PAMAM dendrimers, that can be readily synthesized and have been widely studied for many 

applications. Attachment of various functionalities and drugs on periphery and in the interior 

can also be carried out with benzyl ether dendrimers using recently developed synthetic 

strategies. 47–50 

One of the current goals in PDT is to design porphyrin sensitizers that can absorb in the near 

infra-red (NIR) region. This is highly desirable because skin is more transparent to light of this 

wavelength. However, porphyrins have low two-photon absorption (TPA) cross-sections that limits 

their use in singlet oxygen generation. Direct chemical modification of the macrocycles has been 

attempted to enhance the TPA cross-section and water solubility. 21 An alternate strategy using 

dendrimers was recently developed where donor chromophores with a high TPA cross-section were 

covalently attached to the periphery of a benzyl ether dendrimer, containing a porphyrin moiety at 

the core. 51 It was observed that emission from porphyrin moiety in the dendritic derivative was 17 

times higher than from a porphyrin without the dendritic donor substituents. These dendrimers were 

then decorated with water-soluble triethyleneglycol moieties to further expand the scope of their 

application. 52 In another study, benzyl ether dendrimers with porphyrin unit at the core were made 

water-soluble by attachment of quaternary ammonium or carboxylic groups (Figure 26.5), and their 

efficiency as photosensitizers was evaluated. 53 To further enhance water-solubility and cellular 

uptake of the dendrimer porphyrins with charged periphery, poly-ion complex (PIC) micelles 

composed of polyethylene glycol–poly(L-lysine) block copolymers were developed that could 

entrap the negatively charged dendrimers through electrostatic association (Figure 26.6). 54 These 

polymeric micelle-based nanocarriers showed enhanced cellular uptake, higher photocytotoxicity, 

and similar singlet oxygen generation compared to free dendrimer porphyrins. The higher cellular 

uptake was due to shielding of negatively charged dendrimer periphery inside a charge-neutral 

micelle. The PIC micelles carrying anionic dendrimers were further tested in vivo for treatment of 

age-related macular degeneration (AMD) caused by choroidal neovascularization (CNV). 55 Experimental 

CNV lesions were created in rats by laser photocoagulation, and PIC micelles were injected 

after seven days. The dendrimer-loaded micelles were found to accumulate preferentially in CNV 

lesions and were retained even after 24 h, whereas free dendrimers were cleared within 24 h. When 

cationic dendrimer porphyrins were encapsulated in the polyethylene glycol–poly(aspartic acid) 

block copolymers, the cellular uptake was lower, and the photodynamic efficacy was higher 

compared to free dendrimers. 56 In all these PIC micelles, the dark systemic toxicity (cytotoxicity 

of dendrimers under dark conditions) was found to be reduced probably as a result of a biocompatible 

poly(ethylene glycol) (PEG) shell. 

In a different approach to PDT, 5-aminolevulinic acid (ALA), a natural precursor of protoporphyrin 

IX, was attached to the periphery of first and second generation dendrimers via ester 

linkages to obtain dendrimer–ALA prodrugs (Figure 26.7). 57 ALA can be intracellularly released 

by non-specific esterases leading to higher intracellular ALA levels and higher protoporphyrin IX 

concentration. Also, the dendrimer–ALA conjugate is more lipophilic than only ALA and can 

easily cross the cell membranes as was evident by higher uptake of the second-generation 

dendrimer studied using tumorigenic cell line PAM 212 in vitro. These dendrimers also showed 

negligible dark systemic toxicity. 

26.3 ENCAPSULATION OF THE DRUG IN A DENDRIMER 


Most of the anti-cancer drugs are hydrophobic, and their entrapment in a hydrophilic dendrimer is a 

useful strategy to enhance their solubility for drug delivery. This kind of drug carriers composed of 

dendrimeric micelles have been a subject of a recent review 12 and is discussed here only briefly. 



X 

X 

X 

X 

X 

X 

X 

X 

O 

O 

O 

O 

O 

O 

X 

X 

O 

O 

X 

O 

O 

O 

X 

O 

O 

O 

O 

O 

O 

O 

O 

O 

X 

O 

X 

O 

O 

O 

O 

O 

N 

X 

O 

X 

O 

N 

N 

O O 

Physical encapsulation prevents the drug from being prematurely degraded. The periphery of the 

dendrimer is rendered hydrophilic, either by attachment of oligoethylene glycol chains or charged 

functional groups such as carboxylates, and the hydrophobic interior is suitable for sequestration of 

hydrophobic molecules. 58–61 The drug loading capacity in a dendrimer by encapsulation depends 

upon its generation and size; however, higher payloads can be obtained only by attaching the drug 

molecules to the peripheral groups. Several examples of modified PAMAM dendrimers used for 

physical encapsulation of drugs have been reported. 62,63 

Melamine-based dendrimers modified on the surface with cationic, anionic, and oligoethylene 

glycol groups have been used to encapsulate methotrexate and 6-mercaptopurine, and their hepatotoxicity 

has been studied in vivo. 64,65 A synthetic strategy has been developed for these 

dendrimers as well to incorporate multiple functionalities on the periphery. A hydrophobic anticancer 

drug, 10-hydroxycamptothecin, was encapsulated in hydrophilic poly(glycerol succinic 

acid) (PGLSA) dendrimers. 66 Cytotoxicity assays of these drug-loaded dendrimers on human 

breast cancer cells showed that the drug retained its activity upon encapsulation. 

N 

O 

O 

O 

O 

O 

X X 

FIGURE 26.5 A water-soluble benzyl ether dendrimer with porphyrin moiety at the core for PDT. 


Zn 

X 

O 

O 

X 

O 

O 

O 

O 

O 

O 

O 

X 

O 

O 

O 

O 

X 

O 

O 

X 

O 

O 

O 

O 

X 

O 

O 

X 

X 

X 

X 

X 

X 

X 

X 

X = COOH

516 

x 

x 

x 

x 

x 

x 

x 

x 

O 

O 

O 

O 

O 

O 

x 

O 

O 

x 

O 

x 

O 

O O 

x 

O 

O 

O 

O 

O 

O 

x 

O 

O 

O 

O 

O 

O 

O 

O 

x 

O O 

H 

N Zn N 

H 

x x 

O 

O 

O 

O 

O 

O 

O O O 

O 

x x x x 

O 

O O 

O 

O 

O 

Negatively charged dendrimer 

with a porphyrin core 

O 

O 

O 

O 

O 

x 

O 

O 

x 

O 

x 

O 

O 

O 

O 

x 

O 

O 

H 

N 

x 

x 

x 

x 

x 

x 

x 

x 

O 

O 

O 

O 

O 

O 

O 

NH 

NH 

O 

O 

O 

O 

O 

O 

O 

PEG-b-PLL 

HN 

O 

N 

H 

O 

O 

HN 

O 

O 

O 

H 

N 

NH 

O NH 

FIGURE 26.7 Boc-protected ALA–dendrimer conjugate for PDT. 

O 

O 

HN 

O 

O 

HN 

O 

O 

O 

O 

O 

HN 

O 

O 

O 

O 

Poly-ion complex (PIC) micelle 

FIGURE 26.6 A cartoon showing poly-ion complex (PIC) micelles encapsulating dendrimer–porphyrins. 



O 

O 

O 

O 

O 

O 

O HN 

O 

NH 

O 

O 

O 

O 

NH 

O 

O 

O


26.4 POLYMERIC DRUG CARRIERS BASED ON DENDRIMERS 

Dendrimers have been explored as building blocks in the construction of macromolecular scaffolds 

for drug delivery to circumvent the following problems: rapid release of drugs physically encapsulated 

in a dendrimer, and shorter circulation time of dendrimers themselves. The dendrimer 

component in linear–dendritic hybrid polymers provides the large number of functionalities 

required for high payload, whereas the linear component affords high molecular weight unattainable 

in dendrimers. Such a combination of desirable properties in these hybrid polymers is expected 

to provide higher plasma half-life (circulation time) and enhanced cellular uptake because of EPR 

effect and high potency. With this premise, dendritic aliphatic polyesters to which doxorubicin 

(DOX) was attached via an acid-labile hydrazone linker were connected to a three-arm polyethylene 

glycol star polymer to obtain high molecular weight hybrid polymers (Figure 26.8). 67 This 

polymer–drug conjugate showed less cytotoxicity (human breast cancer cell lines MDA-MD-231 

and MDA-MD-435 using sulforhodamine B assay) than the free drug and longer circulation time 

(tested using 125 I radiolabeling) than only dendrimers. These dendritic aliphatic polyesters were 

also used to prepare linear–dendritic bow-tie hybrids by attaching PEG chains to dendrimer periphery 

(illustrated by a schematic in Figure 26.9) and evaluated for biodistribution. 68 Bow-tie polymers 

with molecular weights up to 1,60,000 g/mol were synthesized, and those with higher molecular 

weight (O40,000 g/mol) and more number of linear arms were studied. When tested on mice 

bearing subcutaneous B16F10 tumors, these polymers were found to accumulate mainly in 

R 

OH 

R 

N 

R 

OH 

N 

OH 

NH 

N 

O 

NH 

O 

O 

R 

NH 

O 

O 

O 

O 

N 

OH 

O 

NH 

O 

O 

O 

O 

O 

O 

HO 

O 

n 

O 

O 

O O O 

O 

n 

O O 

O 

O 

O 

O 

O 

N O 

R 

N 

O 

O 

O O 

O 

NH NH 

HO 

R HO 

N 

N 

R 

R HO 

NH 

NH 

O 

O 

O 

O 

O 

O 

O 

O 

O 

O 

O 

O 

O 

O 

n 

O 

O 

O 

O 

O 

O 

O 

O 

NH 

O 

O 

O 

O 

NH 

R 

NH N 

NH 

N R 

HO 

N 

N 

H2N O 

O 

OH 

CH3 OH O OCH3 OH O 

FIGURE 26.8 A linear–dendritic hybrid star polymer with doxorubicin molecules attached to the periphery 

via acid-labile hydrazone linkages. 


R = 

DOX 

HO 

HO 

OH 

R 

R

518 

tumors with little concentration in the liver, kidney, and lungs. Star-like polymers containing DOX 

have also been synthesized with PAMAM dendrimer as the core. For example, starlike poly-N- 

(2-hydroxypropyl) methacrylamide (HPMA) was synthesized by attachment of the copolymers to 

PAMAM dendrimers of generation 2, 3, and 4. 69 DOX was attached to the methacrylamide backbone 

via a peptide linker (GlyPheLeuGly) because this linker has been shown to get cleaved by an 

enzyme Cathepsin B that is over-expressed in malignant cells. 70 The effect of molecular architecture 

on release of DOX in vitro was clearly evident by the slower release of DOX from the star-like 

polymer compared to that from the linear polymer. 

26.5 SUMMARY 

Dendrimers have clearly emerged as important tools in drug delivery systems. Although tedious 

synthesis of dendrimers hampers their widespread use in biological applications, studies on 

commercially available dendrimers have clearly demonstrated their potential in drug delivery. 

Toward realizing their full potential, however, a combination of useful structural features of 

dendrimers with those of well-established biocompatible linear polymers is necessary. This will 

provide the key to address important issues involving nanoscale medical devices in cancer therapy. 

REFERENCES 

Drug moiety 

Dendrimer 


Linear polymer 

FIGURE 26.9 Structure of a dendritic–linear bow-tie hybrid polymer is illustrated with a cartoon. 

1. Hedrick, J. L. et al., Application of complex macromolecular architectures for advanced microelectronic 

materials, Chem. Eur. J., 8, 3308, 2002. 

2. Son, S. J. et al., Magnetic nanotubes for magnetic-field-assisted bioseparation, biointeraction, and 

drug delivery, J. Am. Chem. Soc., 127, 7316, 2005. 

3. Weissleder, R. et al., Cell-specific targeting of nanoparticles by multivalent attachment of small 

molecules, Nat. Biotechnol., 23, 1418, 2005. 

4. Katz, E. and Willner, I., Nanobiotechnology: Integrated nanoparticle–biomolecule hybrid systems: 

Synthesis, properties, and applications, Angew. Chem. Int. Ed., 43, 6042, 2004. 


6. Gillies, E. R. and Frechet, J. M. J., Dendrimers and dendritic polymers in drug delivery, Drug Discov. 

Today, 10, 35, 2005. 

7. Kim, Y. and Zimmerman, S. C., Applications of dendrimers in bio-organic chemistry, Curr. Opin. 

Chem. Biol., 2, 733, 1998. 



8. Boas, U. and Heegaard, P. M. H., Dendrimers in drug research, Chem. Soc. Rev., 33, 43, 2004. 

9. Tully, D. C. and Frechet, J. M. J., Dendrimers at surfaces and interfaces: Chemistry and applications, 

Chem. Commun., 14, 1229, 2001. 


delivery, Curr. Opin. Chem. Biol., 6, 466, 2002. 

11. Haag, R. and Vogtle, F., Highly branched macromolecules at the interface of chemistry, biology, 

physics, and medicine, Angew. Chem. Int. Ed., 43, 272, 2004. 


release and targeted delivery, Mol. Pharm., 2, 264, 2005. 

13. Bosman, A. W., Janssen, H. M., and Meijer, E. W., About dendrimers: Structure, physical properties, 

and applications, Chem. Rev., 99, 1665, 1999. 

14. Basu, S., Sandanaraj, B. S., and Thayumanavan, S., Molecular recognition in dendrimers, In Encyclopedia 

of Polymer Science and Technology, 4th ed., Mark, H. F., Ed., Wiley, New York, 2004. 

15. Grayson, S. M. and Frechet, J. M. J., Convergent dendrons and dendrimers: From synthesis to 

applications, Chem. Rev., 101, 3819, 2001. 

16. Newkome, G. R., Moorefield, C. N., and Vogtle, F., Eds., Dendrimers and Dendrons: Concepts, 

Syntheses, Applications, 2nd ed., Wiley-VCH, Weinheim, 2001. 

17. Abu-Rmaileh, R., Attwood, D., and D’Emanuele, A., Dendrimers in cancer therapy, Drug Deliv. Syst. 

Sci., 3, 65, 2003. 

18. Sampathkumar, S. -G. and Yarema, K. J., Targeting cancer cells with dendrimers, Chem. Biol., 12, 5, 

2005. 


chemotherapy: Mechanism of tumoritropic accumulation of proteins and antitumor agent smancs, 

Cancer Res., 46, 6387, 1986. 

20. Henderson, B. W. and Dougherty, T. J., Eds., Photodynamic Therapy: Basic Principles and Clinical 

Applications, Marcel Dekker, New York, 1992. 

21. Lunardi, C. N. and Tedesco, A. C., Synergic photosensitizers: A new trend in photodynamic therapy, 

Curr. Org. Chem., 9, 813, 2005. 

22. Barth, R. F. et al., Boron neutron capture therapy of cancer: Current status and future prospects, Clin. 

Cancer Res., 11, 3987, 2005. 

23. Matthews, S. E., Pouton, C. W., and Threadgill, M. D., Macromolecular systems for chemotherapy 

and magnetic resonance imaging, Adv. Drug Deliv. Rev., 18, 219, 1996. 

24. De Groot, F. M. H. et al., Cascade-release dendrimers liberate all end groups upon a single triggering 

event in the dendritic core, Angew. Chem. Int. Ed., 42, 4490, 2003. 

25. Esfand, R. and Tomalia, D. A., Poly(amidoamine) (PAMAM) dendrimers: From biomimicry to drug 

delivery and biomedical applications, Drug Discov. Today, 6, 427, 2001. 

26. Malik, N. et al., Dendrimers: Relationship between structure and biocompatibility in vitro, and 

preliminary studies on the biodistribution of 125 I labeled polyamidoamine dendrimers in vivo, 


27. Shi, X., Majoros, I. J., and Baker, J. R., Capillary electrophoresis of poly(amidoamine) dendrimers: 

From simple derivatives to complex multifunctional medical nanodevices, Mol. Pharm., 2, 278, 2005. 

28. Khandare, J. et al., Synthesis, cellular transport and activity of polyamidoamine dendrimer-methylprednisolone 

conjugates, Bioconjug. Chem., 16, 330, 2005. 

29. Patri, A. K. et al., Synthesis and in vitro testing of J591 antibody–dendrimer conjugate for targeted 

prostate cancer therapy, Bioconjug. Chem., 15, 1174, 2004. 

30. Majoros, I. J. et al., Poly(amidoamine) dendrimer-based multifunctional engineered nanodevice for 

cancer therapy, J. Med. Chem., 48, 5892, 2005. 

31. Choi, Y. et al., Synthesis and functional evaluation of DNA-assembled polyamidoamine dendrimer 

clusters for cancer cell-specific targeting, Chem. Biol., 12, 35, 2005. 

32. Bezouska, K., Design, functional evaluation and biomedical applications of carbohydrate dendrimers 

(glycodendrimers), Rev. Mol. Biotechnol., 90, 269, 2002. 

33. Woller, E. K. et al., Altering the strength of lectin binding interactions and controlling the amount of 

lectin clustering using mannose/hydroxyl-functionalized dendrimers, J. Am. Chem. Soc., 125, 8820, 

2003. 


520 


34. Page, D. and Roy, R., Synthesis and biological properties of mannosylated starburst poly(amidoamine) 

dendrimers, Bioconjug. Chem., 8, 714, 1997. 

35. Venditto, V. J. et al., PAMAM dendrimer based macromolecules as improved contrast agents, Mol. 

Pharm., 2, 302, 2005. 

36. Kobayashi, H. et al., Avidin–dendrimer–(1B4M-Gd)254: A tumor-targeting therapeutic agent for 

gadolinium neutron capture therapy of intraperitoneal disseminated tumor which can be monitored 

by MRI, Bioconjug. Chem., 12, 587, 2001. 

37. Wiener, E. and Narayanan, V. V., Magnetic resonance imaging contrast agents: Theory and the role of 

dendrimers, Adv. Dendr. Macromol., 5, 129, 2002. 

38. Stiriba, S. E., Frey, H., and Haag, R., Dendritic polymers in biomedical applications: From potential to 

clinical use in diagnostics and therapy, Angew. Chem. Int. Ed., 41, 1329, 2002. 

39. Kobayashi, H. et al., Rapid accumulation and internalization of radiolabeled herceptin in an inflammatory 

breast cancer xenograft with vasculogenic mimicry predicted by the contrast-enhanced 

dynamic MRI with the macromolecular contrast agent G6-(1B4M-Gd)256, Cancer Res., 62, 860, 2002. 


dendrimers as potential agents for neutron capture therapy, Bioconjug. Chem., 14, 158, 2003. 


capture therapy of brain tumors, Bioconjug. Chem., 7, 7, 1996. 

42. Kobayashi, H. et al., Activated clearance of a biotinylated macromolecular MRI contrast agent from 

the blood pool using an avidin chase, Bioconjug. Chem., 14, 1044, 2003. 

43. Pierre, V. C., Botta, M., and Raymond, K. N., Dendrimeric gadolinium chelate with fast water 

exchange and high relaxivity at high magnetic field strength, J. Am. Chem. Soc., 127, 504, 2005. 

44. Newkome, G. R. et al., Chemistry of micelles. 37. Internal chemical transformations in a precursor of a 

unimolecular micelle: Boron supercluster via site-specific addition of B10H14 to cascade molecules, 

Angew. Chem. Int. Ed. Engl., 33, 666, 1994. 

45. Parrott, M. C. et al., Synthesis and properties of carborane-functionalized aliphatic polyester dendrimers, 

J. Am. Chem. Soc., 127, 12081, 2005. 

46. Brewis, M. et al., Phthalocyanine-centred aryl ether dendrimers with oligo(ethyleneoxy) surface 

groups, Tetrahedron Lett., 42, 813, 2001. 

47. Freeman, A. W., Chrisstoffels, L. A. J., and Frechet, J. M. J., A Simple method for controlling 

dendritic architecture and diversity: A parallel monomer combination approach, J. Org. Chem., 65, 

7612, 2000. 

48. Sivanandan, K., Sandanaraj, B. S., and Thayumanavan, S., Sequences in dendrons and dendrimers, 

J. Org. Chem., 69, 2937, 2004. 

49. Vutukuri, D., Sivanandan, K., and Thayumanavan, S., Synthesis of dendrimers with multifunctional 

periphery using an ABB 0 monomer, Chem. Commun., 21, 796–797, 2003. 

50. Sivanandan, K., Vutukuri, D., and Thayumanavan, S., Functional group diversity in dendrimers, Org. 

Lett., 4, 3751, 2002. 

51. Dichtel, W. R. et al., Singlet oxygen generation via two-photon excited FRET, J. Am. Chem. Soc., 126, 

5380, 2004. 

52. Oar, M. A. et al., Photosensitization of singlet oxygen via two-photon excited fluorescence resonance 

energy transfer in a water-soluble dendrimer, Chem. Mater., 17, 2267, 2005. 

53. Nishiyama, N. et al., Light-harvesting ionic dendrimer porphyrins as new photosensitizers for photodynamic 

therapy, Bioconjug. Chem., 14, 58, 2003. 

54. Jang, W. -D. et al., Supramolecular nanocarrier of anionic dendrimer porphyrins with cationic block 

copolymers modified with polyethylene glycol to enhance intracellular photodynamic efficacy, 

Angew. Chem. Int. Ed., 44, 419, 2005. 

55. Ideta, R. et al., Nanotechnology-based photodynamic therapy for neovascular disease using a supramolecular 

nanocarrier loaded with a dendritic photosensitizer, Nano Lett., 5(12), 2426–2431, 2005. 

56. Zhang, G. -D. et al., Polyion complex micelles entrapping cationic dendrimer porphyrin: Effective 

photosensitizer for photodynamic therapy of cancer, J. Control Release, 93, 141, 2003. 

57. Battah, S. H. et al., Synthesis and biological studies of 5-aminolevulinic acid-containing dendrimers 

for photodynamic therapy, Bioconjug. Chem., 12, 980, 2001. 

58. Vutukuri, D. R., Basu, S., and Thayumanavan, S., Dendrimers with both polar and apolar nanocontainer 

characteristics, J. Am. Chem. Soc., 126, 15636, 2004. 



59. Gopidas, K. R., Whitesell, J. K., and Fox, M. A., Metal-core-organic shell dendrimers as unimolecular 

micelles, J. Am. Chem. Soc., 125, 14168, 2003. 

60. Hawker, C. J., Wooley, K. L., and Frechet, J. M. J., Unimolecular micelles and globular amphiphiles: 

Dendritic macromolecules as novel recyclable solubilization agents, J. Chem. Soc. Perkin Trans., 1, 

1287, 1993. 

61. Aathimanikandan, S. V., Savariar, E. N., and Thayumanavan, S., Temperature-sensitive dendritic 

micelles, J. Am. Chem. Soc., 127, 14922, 2005. 

62. Beezer, A. E. et al., Dendrimers as potential drug carriers; encapsulation of acidic hydrophobes within 

water soluble PAMAM derivatives, Tetrahedron, 59, 3873, 2003. 


their ability to encapsulate anticancer drugs, Bioconjug. Chem., 11, 910, 2000. 

64. Neerman, M. F. et al., Reduction of drug toxicity using dendrimers based on melamine, Mol. Pharm., 

1, 390, 2004. 

65. Zhang, W. et al., Orthogonal, convergent syntheses of dendrimers based on melamine with one or two 

unique surface sites for manipulation, J. Am. Chem. Soc., 123, 8914, 2001. 

66. Morgan, M. T. et al., Dendritic molecular capsules for hydrophobic compounds, J. Am. Chem. Soc., 

125, 15485, 2003. 

67. Padilla De Jesus, O. L. et al., Polyester dendritic systems for drug delivery applications: In vitro and 

in vivo evaluation, Bioconjug. Chem., 13, 453, 2002. 

68. Gillies, E. R. et al., Biological evaluation of polyester dendrimer: Poly(ethylene oxide) “bow-tie” 

hybrids with tunable molecular weight and architecture, Mol. Pharm., 2, 129, 2005. 

69. Wang, D. et al., Synthesis of starlike N-(2-hydroxypropyl) methacrylamide copolymers: Potential 

drug carriers, Biomacromolecules, 1, 313, 2000. 

70. Fuchs, S. et al., Fluorescent dendrimers with a peptide cathepsin B cleavage site for drug delivery 

applications, Chem. Commun., 14, 1830–1832, 2005. 


27 

CONTENTS 

PEGylated Dendritic 

Nanoparticulate Carriers of 

Anti-Cancer Drugs 

D. Bhadra, S. Bhadra, and N. K. Jain 

27.1 Introduction ....................................................................................................................... 523 

27.2 Problems in Cancer Chemotherapy .................................................................................. 524 

27.3 Nanoparticles in Cancer Chemotherapy ........................................................................... 525 

27.3.1 Challenges of Nanotechnology ........................................................................... 527 

27.4 Poly(Ethylene Glycol) Conjugation .................................................................................. 527 

27.4.1 Derivatization and Activation of Poly(Ethylene Glycol) ................................... 528 

27.4.2 Properties of Poly(Ethylene Glycol) Conjugates................................................ 528 

27.5 Dendrimer as Drug Delivery System: Selection, Applicability, and Rationality ............ 529 

27.5.1 Differences Between Hyperbranched Structures and Dendrimers ..................... 532 

27.5.2 Classification of Dendrimers............................................................................... 532 

27.5.3 Features of Dendrimers as Nanodrugs................................................................ 533 

27.6 Applications of PEGylated Dendrimers............................................................................ 534 

27.6.1 Enhanced Biocompatibility for Drug Delivery .................................................. 535 

27.6.2 Micellar Means of Enhancing Solubility............................................................ 537 

27.6.3 Enhancing Pharmacokinetic Properties .............................................................. 538 

27.6.4 Dendrimer Use as Transfection Reagents .......................................................... 539 

27.6.5 Stimuli Responsive Carriers................................................................................ 540 

27.6.6 Alteration of Toxicity.......................................................................................... 541 

27.6.7 Effective Photodynamic Therapy........................................................................ 543 

27.6.8 Dendritic Gels ..................................................................................................... 543 

27.6.9 Ligand-Based Drug Delivery .............................................................................. 543 

27.6.10 Increasingly Effective Boron Neutron Capture Therapy.................................... 544 

27.6.11 Improved MRI Contrast Agents.......................................................................... 544 

27.7 Conclusions ....................................................................................................................... 545 

References .................................................................................................................................... 545 


Cancer is a leading cause of death. Irrespective of etiology, cancer is basically a disease of cells 

characterized by loss of normal cellular growth, maturation, and multiplication that lead to disturbance 

of homeostasis. The main features of cancer are (Barar 2000) excessive cell growth, 


523

524 

invasiveness, undifferentiated cells or tissues, the ability to metastasize or spread to newer sites and 

establish new growths, a type of acquired heredity where the progeny of cancer cells also retain 

cancerous properties, and a shift of cellular metabolism, leading to increased production of macromolecules 

from nucleosides and amino acids that causes an increased metabolism of carbohydrates 

for cellular energy, ultimately resulting in the host’s illness. This illness is attributed to pressure 

effects as a result of local tumor growth, destruction of the organs involved by the primary growth 

or its metastases, and systemic effects as a result of the new growths. 

A single cancerous cell surrounded by healthy tissue will replicate at a rate higher than the other 

cells, placing a strain on the nutrient supply and elimination of metabolic waste products. Tumor cells 

will displace healthy cells until the tumor reaches a diffusion-limited maximal size. Although tumor 

cells will typically not initiate apoptosis in a low nutrient environment, they do require the normal 

building blocks of cell function like oxygen, glucose, and amino acids. The vasculature that was 

designed to supply the now-extinct healthy tissue could not place as high a demand for nutrients 

because of its slower growth rate. Tumor cells will continue dividing because they do so without 

regard to nutrient supply, but many tumor cells will perish because the amount of nutrients is 

insufficient. The tumor cells at the outer edge of a mass have the best access to nutrients, and cells 

on the inside die creating a necrotic core within tumors that rely on diffusion to deliver nutrients and 

eliminate waste products. In essence, a steady state tumor size forms as the rate of proliferation is 

equal to the rate of cell death until a better connection with the circulatory system is created. This 

diffusion-limited maximal size of most tumors is around 2 mm 3 (Grossfeld, Carrol, and Lindeman 

2002). To grow beyond this size, the tumor must recruit the formation of blood vessels to provide the 

necessary nutrients to fuel its continued expansion. It is thought that there could be numerous tumors 

at this diffusion-limited maximal size throughout the body. Until the tumor can gain that access to the 

circulation, it will remain at this size, and the process can take years. 

27.2 PROBLEMS IN CANCER CHEMOTHERAPY 


In the past, chemotherapy was considered as a last resort after more successful treatments like 

surgery and radiotherapy had failed. The main problem of cancer chemotherapy is the lack of 

highly selective drugs, and the rapidly dividing normal cells of the bone marrow, gut, lymphoid 

tissue, spermatogenic cells, fetus, and hair follicles are also killed. Most of the antineoplastic drugs 

act on the processes such as DNA synthesis, transcription, or the mitotic phase, and they are labeled 

as cell cycle phase-specific drugs (also known as phase-dependent drugs). The phase-specific drugs 

do not act on G 0 phase. In contrast, there are certain drugs that kill the cells during all or most 

phases of the cycle, and they are labeled as cell cycle phase-nonspecific drugs (also known as 

phase-independent drugs). The phase-specific drugs have proven to be effective in hematological 

malignancies and tumors with high rate of proliferation or high growth fraction (Barar 2000). 

The various obstacles for drug delivery to tumors include differences in cellular morphology, 

tissue immunogenecity, uncontrolled rate of growth, capacity of metastasize, and poor response to 

chemotherapeutic agents. Furthermore, there are variations in blood flow and vessel permeability 

within different regions of some tissues (Reynolds 1996). Various drug delivery systems are 

employed for delivery of chemotherapeutic agents to neoplastic cells. This helps in targeting 

drugs directly into cells and preventing drug interaction with normal tissues and alleviating side 

effects to normal cells. Because of a lack of highly selective drugs for tumor cells and tissues and 

decreased penetration of drug into neoplastic cells, a number of novel drug delivery systems, 

prodrugs, and other chemically modified forms of drugs having altered tissue distributions were 

brought in use and recently reviewed for targeting to cells located in various parts of body (Brigger, 

Dubernet, and Couvreur 2002). 

There are various means to effectively fight the tumors such as achieving targeting by 

avoiding reticuloendothelial system (RES); targeted delivery through enhanced permeability 


PEGylated Dendritic Nanoparticulate Carriers of Anti-Cancer Drugs 525 

and retention (EPR) effects; tumor-specific targeting; targeting through angiogenesis; targeting 

tumor vasculaturel; etc. (Brannon-Peppas and Blanchette 2004). Particles with longer circulation 

times and greater ability to target to the site of interest should be 100 nm or less in diameter and 

have a hydrophilic surface in order to reduce clearance by macrophages (Storm et al. 1995). 

Coatings of hydrophilic polymers can create a cloud of chains at the particle surface that will 

repel plasma proteins, and work in this area began by adsorbing surfactants to the 

nanoparticles surface. 

Other routes include forming the particles from branched or block copolymers with hydrophilic 

and hydrophobic domains. One potential advantage in treatment of advance stage cancerous tissue 

is the inherent leaky vasculature that allows for greater accumulation of colloidal systems. The 

check the defective vascular architecture, created as a result of the rapid vascularization necessary 

to serve fast-growing cancers, coupled with poor lymphatic drainage allows an enhanced permeation 

and retention effect (EPR effect) (Sledge and Miller 2003). The ability to target 

treatment to very specific cancer cells also uses a cancer’s own structure in that many cancers 

over express particular antigens, even on their surface. This makes them ideal targets for drug 

delivery as long as the targets for a particular cancer cell type can be identified with confidence and 

are not expressed in significant quantities anywhere else in the body. Tumor-activated prodrug 

therapy uses the approach that a drug conjugated to a tumor-specific molecule will remain inactive 

until it reaches the tumor (Chari 1998). These systems would ideally be dependent on interactions 

with cells found specifically on the surface of cancerous cells and not the surface of healthy cells. 

Most linkers are usually peptidase cleavable or acid labile but may not be stable enough in vivo to 

give desirable clinical outcomes. 

Limitations also exist because of the lower potency of some drugs after being linked to 

targeting moieties when the targeting portion is not cleaved correctly or at all. For example, 

adriamycin-conjugated poly(ethylene glycol) linker with enzymatically cleavable peptide 

sequences (alanyl-valine, alanyl-proline, and glycyl-proline) or using monoclonal antibodies has 

shown a greater selectivity to cleavage at tumor cells (Suzawa et al. 2002). A number of targeted 

cancer treatments using antibodies for specific cancer types have been approved by the U.S. Food 

and Drug Administration like rituximab, trastuzumab, gemtuzumabozogamicin, alemtuzumab, 

ibritumomab tiuxetan, gefitinib, etc. (Abou-Jawde et al. 2003). 

27.3 NANOPARTICLES IN CANCER CHEMOTHERAPY 

Nanotechnology applied to cancer treatment may offer several promising advantages over conventional 

drugs. Nanoscale devices are two orders of magnitude smaller than tumor cells, making it 

possible for them to directly interact with intracellular organelles and proteins. Because of their 

molecule-like size, nanoscale tools may be capable of early disease detection using minimal amounts 

of tissue, even down to a single malignant cell (NCI 2005). These tools may not only prevent disease 

by monitoring genetic damage, but also treat cells in vivo while minimizing interference with healthy 

tissue. By combining different kinds of nanoscale tools on a single device, it may be possible to run 

multiple diagnostic tests simultaneously (www.math.uci.edu/~cristini/publications/nanochap.pdf). 

In particular, it is hoped that cancer drug therapy involving nanotechnology will be more effective in 

targeting malignant cells and sparing healthy tissue. In this regard, the role of nanoparticles loaded 

with chemotherapeutic drugs has been receiving much attention. Research and development in this 

area is expected to dramatically increase in importance in the coming years. In general, nanoscale 

drug delivery systems (Figure 27.1) for chemotherapy can be divided into two categories: polymerand 

lipid-based (Langer 2000; Sahoo and Labhasetwar, 2003). 

Polymers are usually larger than lipid molecules, and they form a solid phase such as polymeric 

nanoparticles, films, pellets, dendrimers; whereas, lipids form a liquid (or liquid crystalline phase) 

such as liposomes, cubersomes, micelles, and other emulsions (Feng and Chien 2003). Whereas 


526 

ligand 

Drug 

Drug 

Nanosphere Nanocapsule 

PEG 

ligand 

ligand 

Liposome 

ligand 

ZnS 

CdSe 

ligand 

Antibody 

ligand 

ligand 

ligand 

Ligand conjugated 

ZnS-capped CdSe quantum dots 

Hydrophobic 

core 

Polymeric micelles 

polymer-based systems are considered biologically more stable than lipid-based systems, the latter 

are generally more biocompatible. Polymer-based systems might possess good drug targeting 

ability because their uptake may be different for cells in different tissues (Maruyama 2000). In 

fact, Feng and Chien (2003) have suggested that a combination of polymer- and lipid-based systems 

could integrate their advantages while avoiding their respective disadvantages. An example of such 

a nanoparticle would be a liposomes-in-microspheres (LIM) system where drugs are first loaded 

into liposomes and then encapsulated into polymeric microspheres. This way, both hydrophobic 

and hydrophilic drugs can be delivered in one nanoparticle. The bioactivity of peptides and proteins 

would be preserved in the liposomes whose stability is protected by the polymeric matrix (Feng and 

Chien 2003). 

Chemotherapy using nanoparticles has been studied in clinical trials for several years, and 

numerous studies have been published in this regard (Kreuter 1994). Two liposomally 

delivered drugs are currently on the market: daunorubicin and doxorubicin (Massing and 

Fuxius 2000). These encapsulated drugs can be formulated to maximize their half-life in 

the circulation. For example, a stealth version of liposomal doxorubicin coated with polyethylene 

glycol to reduce its uptake by the RES can extend its half-life in blood for up to 50–60 h 

(CCO Formulary 2003). 

Core 


Internal cavity 

Branchingunits 

Fourth generation dendrimer 

Fe3O 4 

Ligend 

Ferrofluid conjugated to ligand 

FIGURE 27.1 Representation of different nanoparticulate carriers used for drug delivery. (Adapted from Sahoo, 

S. K. and Labhasetwar, V., Drug Discov. Today, 8, 2003, www.drugdiscoverytoday.com. With permission.) 



27.3.1 CHALLENGES OF NANOTECHNOLOGY 

The difficulties faced by nanotechnology in the service of clinical medicine are numerous. These 

difficulties should be kept in mind while considering chemotherapeutic treatment involving nanotechnology 

and the potential role of biocomputation. First, there are basic physical issues that 

matter on a small scale. Because matter behaves differently on the nanolevel than it does at 

micro- and macro levels, most of the science at the nanoscale has been devoted to basic research, 

designed to expand understanding of how matter behaves on this length scale (www.math.uci.edu/~ 

cristini/publications/nanochap.pdf). Because nanomaterials have large surface areas relative to 

their volumes, phenomena such as friction are more critical than they are in larger systems. The 

small size of nanoparticles may result in significant delay or speed-up in their intended actions. 

They may accumulate at unintended sites in the body. They may provoke unexpected immune 

system reactions. Cells may adapt to the nanoparticles, modifying the body’s behavior in unforeseen 

ways (www.math.uci.edu/~cristini/publications/nanochap.pdf). 

The efficacy of nanoparticles may be adversely affected by their interaction with the cellular 

environment. For instance, the RES may clear nanoscale devices, even stealth versions, too rapidly 

for them to be effective because of the tendency of the RES to phagocytose nanoparticles (Kreuter 

1994). Nanoparticles can be taken up by dendritic cells (Elamanchili et al. 2004) and by macrophages 

(Cui, Hsu, and Mumper 2003). RES accumulation of nanoparticles could potentially lead to 

a compromise of the immune system. On the other hand, larger nanoparticles may accumulate in 

larger organs, leading to toxicity. Perhaps the biggest issue of all is that the physically compromised 

tumor vasculature may prevent most of the nanodevices from reaching the target cells by vascular 

transport or diffusion. Alterations in the tumor vasculature may adversely affect the convection of 

the nanodevices in the blood stream (Brigger, Dubernet, and Couvreur 2002). Local cell density and 

other stromal features may hamper drug or nanodevice diffusion through tumoral tissue. 

27.4 POLY(ETHYLENE GLYCOL) CONJUGATION 

Poly(ethylene glycol) is a non-toxic water-soluble polymer that resists recognition by the immune 

system. The term PEG is used for poly(ethylene glycol) that is often referred for polymer chains 

with molecular weight (MWs) below 20,000 Da, and poly(ethylene oxide) (PEO) refers to higher 

MW polymers (Harris et al. 1984; Peppas et al. 1999). These are homogeneous polymers of Mw/ 

Mn!1.1 represented by the formula HO(CH2CH2O)nH and is called as poly(oxy-1, 2-ethanediyl), 

or a-hydro-u-hydroxy-poly(ethylene glycol). It is an additional polymer of ethylene oxide and 

water, where n represents the average number of oxy-ethylene groups. These are generally soluble 

in water, alcohol, acetone, and chloroform, but miscible with glycols and practically insoluble in 

ether. These are also known as Macrogols (Reynolds 1996) that are generally of higher molecular 

weights such as Macrogols of 20,000 Da. The pH of 5% solution is about 4–7.5. 

The aqueous solutions of PEGs can be sterilized by filtration, autoclaving, etc. and are needed 

to be stored in airtight containers. The PEGs are relatively less toxic, but toxicity, if any appears to 

be greatest with macrogols of lower molecular weights. On topical administration, it may cause 

stinging, especially to mucous membrane. Such administration of PEGs may be associated with 

hypersensitivity reactions such as urticaria, and local gastrointestinal discomfort, bloating and 

nausea, abdominal cramps, vomiting and irritation, may result on its use in bowel cleansing 

preparation. Macrogols demonstrate oxidizing ability, leading to incompatibility with activity of 

bacitracin or benzyl-penicillin that may get reduced in macrogol bases. Estimated acceptable daily 

intake may be upto 10-mg/kg body weight (Reynolds 1996). Its presence in aqueous solution has no 

deleterious effect on protein conformation or activity of enzymes. It exhibits rapid clearance from 

the body, and it has been approved for a wide range of biomedical applications. PEG may transfer 

its properties to other molecules when it is covalently bound to that molecule. This can result in 

modification of number of properties of molecules as illustrated later on (Zalipsky 1995). 


528 

PEG has polyether backbone that is inert in biological environment as well as in most chemical 

reaction conditions as in chemical modifications and/or conjugation reactions. The chemical derivatisation 

of end groups of PEG is an essential first step in preparation of bioconjugates. The 

various properties are altered because of polyethylene glycol conjugation that is discussed in the 

following sections. PEG conjugation was tried and found most applicable in case of many nanoparticulate 

systems, using activated PEGs with appropriate spacers. 

27.4.1 DERIVATIZATION AND ACTIVATION OF POLY(ETHYLENE GLYCOL) 

Poly(ethylene glycol) has two equivalent hydroxyl groups, so this could act as potential crosslinking 

agent for any systems to which it is attached. These hydroxyl groups can be attached to 

various bio-active species by covalent coupling using number of simple PEG analogues. Some 

important stable analogues are bromo, amino, aminoethyl, carboxymethyl, succinimidosuccinate, 

tosylate, mesylate, aldehyde, octadecylamine, monopalmitate, stearoyloxy derivative of PEG, or 

monomethoxy-PEG (MPEG) (Buckmann and Morr 1981; Harris et al. 1984). These were prepared 

for reactions with sensitive bio-active systems under mild conditions. 

One very important derivative that has been used in number of derivatisation reactions that has 

one hydroxyl group blocked is monomethyl-ether of polyethylene-glycol or monomethoxy-polyethylene-glycol 

(MPEG), i.e., considered equivalent in terms of reaction for formation of PEG 

derivatives. This was generally brought in use for conjugation to bioactive species. It is generally 

used when multiple chains of polymers had to be linked to the intended substrates. Because of 

structural simplicity and possession of only one-derivatizable end group, the use of MPEG minimizes 

cross-linking possibilities and leads to improved homogeneity of conjugate. Therefore, it 

usually is a starting material of choice for the covalent modification of proteins, biomaterials, 

particulates, lipids, drugs, etc. (Zalipsky 1995). 

MeOH CHðOCH 2CH 2Þ nOH/MeOðCH 2CH 2OÞ nH 

Commercially available MPEGs are often contaminated with significant amounts (equivalent to 

25%) of dihydroxy-terminated polymers and may be purified by chromatography and ethyl ether 

precipitation. The use of these polymers as starting materials of choice for covalent modification of 

proteins, biomaterials, and particles is dependent for not forming cross-linked conjugates (Harris 

et al. 1984). 

In most cases, commercially available MPEGs of 2000–5000 MWs are used for preparation of 

number of reagents. MPEG-linked and based electrophiles were synthesized and used for linking 

with number of available attachment sites, e.g., amino acids and other nucleophilic groups. These 

are called activated PEGs (Zalipsky 1995). The intermediate linkers are called spacers, linking to 

bioactive groups, especially proteins and enzymes. These generally require mild conditions of 

reactions and specific properties of reactants for binding. 

In order of reaction and based on ease of introduction, these can roughly be arranged as follows: 

arylating agents O acylating agents O alkylating agents (Zalipsky 1995). For example, tresylates, 

succinimidyl succinate derivatives, succinimidyl ester of carboxymethyl PEG, oxy-carboxylamino 

acid derivative of PEG, acetaldehyde, and hydrazide derivatives were mainly used for PEGylation 

of proteins and enzymes. 

27.4.2 PROPERTIES OF POLY(ETHYLENE GLYCOL) CONJUGATES 


This group reviewed various properties affected by PEGylation of carriers (Bhadra et al. 2002). 

PEGylation has pronounced effects on various parameters of drug delivery systems (Katre 1993; 

Zalipsky 1995). As far as bio-distribution and pharmacokinetic, PEG conjugation increases blood 

circulation half-life, by reducing the tissue distribution, RES, liver, spleen, and macrophage uptake 



of the carriers. (Veronese et al. 1989; Papahadjopoulus et al. 1991; Phillips et al. 1999). PEG 

coatings alter solubility of the systems (Choi et al. 1998;Kim et al. 1999). It reduces the toxicity 

by decreasing the release and contact of active constituent, decreasing immunogenicity, celladherence, 

thrombogenicity, and protein adsorption. PEGylation can also lead to stabilization of 

delivery systems. It can also improve utilization as drug a delivery system by increasing drug 

loading per system, improving drug targeting as in case of liposomes by diffusion controlling 

mechanisms, and promoting its use as a micellar delivery system. This process can sustain and 

control drug delivery safely and appropriately by decreasing burst effects and by decreasing drug 

leakage and increasing stability of system. It can improve transfection capability (Choi et al. 1998). 

This can increase tumor uptake (Papahadjopoulus et al. 1991; Ishida et al. 1999) of drug and 

delivery systems. 

Manipulation of the upper layers (corona) of nanoparticles or dendrimers confers advantageous 

properties to such particles, e.g., increased solubility and biocompatibility (McNeil 

2005). Attaching hydrophilic polymers to the surface such as PEG greatly increases the hydration 

(i.e., solubility) of the nanoparticles and can protect attached proteins from enzymatic degradation 

when used for in vivo applications (Bhadra et al. 2002; Harris and Chess 2003). Nanoparticles 

with hydrophilic polymers such as PEG attached to their surface can act as a platform for 

lipophilic molecules, and they can overcome the solubility barrier. Insoluble compounds can 

be attached, adsorbed, or otherwise encapsulated in the hydrated nanoparticles. The surface 

addition of PEG and other hydrophilic polymers also increases the in vivo compatibility of 

nanoparticles. When intravascularly injected, uncoated nanoparticles are rapidly cleared from 

the bloodstream by the reticuloendothelial system (RES) (Brigger, Dubernet, and Couvreur 

2002). Nanoparticles coated with hydrophilic polymers have prolonged half-lives, believed to 

result from decreased opsonization and subsequent clearance by macrophages (Moghimi and 

Szebeni 2003). 

Cancer therapy has benefited from the use of liposomal doxorubicin, a formulation that again 

increases the therapeutic index of the active agent through a combination of passive tumor targeting 

and reduced toxicity (Gabizon and Martin 1997). In this case, coating the liposome with PEG 

significantly decreases uptake by macrophages and allows the liposomes to concentrate in tumors 

by escaping from the leaky vasculature surrounding solid tumors (Dvorak et al. 1988) through EPR 

effect (Maeda 2001). The surface chemistry of the nanoshells can be modified with PEG to increase 

biocompatibility and with sulfide-based linkers to allow the particles to be functionalized with 

targeting ligands. 

27.5 DENDRIMER AS DRUG DELIVERY SYSTEM: SELECTION, 

APPLICABILITY, AND RATIONALITY 

Dendrimers are a relatively new class of polymers with structures that depart rather dramatically 

from traditional linear polymers that are built from so-called AB monomers (Newkome, Moorefield, 

and Vogtle 1996). Because they are built from ABn, monomers (where n is usually 2 or 3), 

dendrimers are highly branched and have three distinct structural features: a core, multiple peripheral 

(end-) groups, and branching units that link the two. The peripheral groups and branching units 

are together called dendrons, or informally, wedges. Branching units used to date have contained 

virtually every type of functional group, ranging from those comprising pure hydrocarbons or 

aromatic groups to modified carbohydrates and nucleic acids. Dendrimers are iteratively synthesized 

so that the number of intervening branching units between the core and one end-group 

(i.e., the number of layers), called the generation number, is determined by the number of synthetic 

cycles. In the commercially available poly(amidoamine) (PAMAM) dendrimers, ammonia represents 

the zeroth generation (Tomalia, Naylor, and Goddard 1990). As the generation number 

increases from one to seven, the number of peripheral groups follows the geometric series 3, 6, 


530 


12,.,192. These peripheral groups largely control the solubility of the compound so that even with 

highly hydrophobic internal units, a dendrimer will dissolve in water if the end groups are sufficiently 

hydrophilic. As will be discussed here, this feature and the high local concentration of end 

groups enable a number of applications in bio-organic chemistry. Dendrimers represent a new class 

of highly branched polymers whose interior cavities and multiple peripheral groups facilitate 

potential applications in biomedicine and bio-organic chemistry (Kim and Zimmerman 1998). 

Major advances in the past years were made in the synthesis and study of new carbohydrate, 

nucleic acid, and peptide dendrimers as well as in the use of dendrimers as magnetic resonance 

imaging contrast agents, agents for cellular delivery of nucleic acids, and scaffolds for 

biomimetic systems. 

Researchers are now trying to synthesize dendrimer clusters for targeted therapy. Scientists at 

the University of Michigan have devised a method to easily create multipurpose molecules for use 

in targeted anti-cancer therapy (NCI 2005). By attaching single-stranded DNA linkers to dendrimers, 

the researchers could bridge together individual dendrimer subunits with specific functions 

and create a multifunctional dendrimer cluster. This bridging technique could potentially lead the 

way to developing customized therapies for individual patients. Dendrimers are promising candidates 

for use as the base of anti-cancer drugs to which functional molecules can be attached. 

Ideally, an anti-cancer therapeutic agent would have multiple functional groups such as a 

targeting molecule to bind a cell, a radioactive compound to kill the cell, and a fluorescent 

probe so the process can be imaged. However, there are problems synthesizing these compounds, 

and different drugs would have to be designed for each different tumor with this approach. 

Dr. James Baker, Jr. and colleagues improved compound synthesis through a building block 

approach in one NCI-funded study. They created two single-function dendrimers; one with a 

molecule to bind folate receptors and the other with a fluorescein molecule for imaging. They 

then added complimentary 34-base stretches of single-stranded DNA to each, and the two dendrimers 

conjugated through base pairing. Using fluorescence imaging, they observed that the 

dendrimer cluster specifically bound to a cancer cell line over expressing the folate receptor. 

Compared with traditional chemistry, DNA-linked dendrimers could provide a more effective 

way to mix and match different therapeutic combinations to better treat individual patients. 

Dendrimers represent one nanostructured material that may soon find its way into medical 

therapeutics. Starburst dendrimers are tree-shaped synthetic molecules with a regular branching 

structure emanating outward from a core that forms nanometer by nanometer with the number of 

synthetic steps or generations dictating the exact size of the particles, typically a few nanometers in 

spheroidal diameter (Freitas 2005). The peripheral layer can be made to form a dense field of 

molecular groups that serve as hooks for attaching other useful molecules such as DNA that can 

enter cells while avoiding triggering an immune response unlike viral vectors commonly employed 

today for transfection. Upon encountering a living cell, dendrimers of a certain size trigger a 

process called endocytosis where the cell’s outermost membrane deforms into a tiny bubble or 

vesicle. The vesicle encloses the dendrimer that is then admitted into the cell’s interior. Once inside, 

the DNA is released and migrates to the nucleus where it becomes part of the cell’s genome. The 

technique has been tested on a variety of mammalian cell types (Kukowska-Latallo et al. 2000) and 

in animal models (Vincent et al. 2003) though clinical human trials of dendrimer gene therapy 

remain to be done. 

Glycodendrimer nanodecoys have also been used to trap and deactivate some strains of influenza 

virus particles (Reuter et al. 1999; Landers et al. 2002). James Baker’s group at the University 

of Michigan is extending this work to the synthesis of multi-component nanodevices called tectodendrimers 

built up from a number of single-molecule dendrimer components (http://nano.med.umich.edu/projects/Dendrimers.html; 

Quintana et al. 2002). Tecto-dendrimers have a single core 

dendrimer surrounded by additional dendrimer modules of different types, each type designed to 

perform a function necessary to a smart therapeutic nanodevice (Figure 27.2). Baker’s group has 



Targeting 

Molecule 

Water 

10 –1 

Liposome 

Corona 

Glucose Antibody Virus Bacteria Cancer Cell A Period Tennis Ball 

Core 

Image Contrast 

Agent 

1 10 10 2 

10 3 

built a library of dendrimeric components from which a combinatorially large number of nanodevices 

can be synthesized. 

The initial library contains components that will perform the following tasks: diseased cell 

recognition, diagnosis of disease state, drug delivery, report location, and report outcome of 

therapy. By using this modular architecture, an array of smart therapeutic nanodevices can be 

created with little effort. For instance, once apoptosis-reporting, contrast-enhancing, and chemotherapeutic-releasing 

dendrimer modules are made and attached to the core dendrimer, it should be 

possible to make large quantities of this tecto-dendrimer as a starting material. This framework 

structure can be customized to fight a particular cancer simply by substituting any one of many 

possible distinct cancer recognition or targeting dendrimers, creating a nanodevice customized to 

destroy a specific cancer type and no other while also sparing the healthy normal cells. These 

nanodevices were synthesized using an ethylenediamine core with folic acid, fluorescein, and 

methotrexate covalently attached to the surface to provide targeting, imaging, and intracellular 

drug delivery capabilities (Freitas 1999). The “targeted delivery improved the cytotoxic response of 

the cells to methotrexate 100-fold over free drug” (Quintana et al. 2002). 

At least a half dozen cancer cell types have already been associated with at least one unique 

protein that targeting dendrimers could use to identify the cell as cancerous, and as the 

10 4 

Dendrimer Gold Nanoshell 

Drugs 

PEG 

(polyethylene 

glycol) 

Research 

Drug Screening 

(Labeling) 

Gene Delivery 

(Transfection) 

Diagnosis 

(Devices and 

Labeling) 

10 5 

10 6 

Quantum Dot 

Nanotechnology 

10 7 

Medical Applications 

10 8 

Nanometers 

Fullerene 

Clinical 

Drug Delivery 

(Therapy) 

Detection 

(Imaging) 

Diagnosis/Monitoring 

(Disease Markers) 

FIGURE 27.2 Morphological and size-based comparison of various novel drug delivery carriers such as 

liposome, dendrimer, gold nanoshell, quantum dot, fullerene, and nanometric scale of representation of 

various similar systems. (Adapted from McNeil, S. E., J. Leukoc. Biol., 78, 1–10, 2005. With permission.) 


532 

genomic revolution progresses, it is likely that proteins unique to each kind of cancer will 

be identified, allowing Baker to design a recognition dendrimer for each type of cancer (http://nano.med.umich.edu/projects/Dendrimers.html). 

The same cell-surface protein recognition-targeting 

strategy could be applied against virus-infected cells and parasites. Molecular modeling has been 

used to determine optimal dendrimer surface modifications for the function of tecto-dendrimer 

nanodevices and to suggest surface modifications that improve targeting (Quintana et al. 2002). 

NASA and the National Cancer Institute have funded Baker’s lab to produce dendrimer-based 

nanodevices that can detect and report cellular damage because of radiation exposure in astronauts 

on long-term space missions (Sparks 2002). By mid-2002, the lab had built a dendrimeric nanodevice 

to detect and report the intracellular presence of caspase-3, one of the first enzymes released 

during cellular suicide or apoptosis (programmed cell death) that is one sign of a radiation-damaged 

cell. The device includes one component that identifies the dendrimer as a blood sugar so that the 

nanodevice is readily absorbed into a white blood cell and a second component using fluorescence 

resonance energy transfer (FRET) that employs two closely bonded molecules. Before apoptosis, 

the FRET system stays bound together, and the white cell interior remains dark upon illumination. 

Once apoptosis begins and caspase-3 is released, the bond is quickly broken, and the white blood 

cell is awash in fluorescent light. If a retinal scanning device measuring the level of fluorescence 

inside an astronaut’s body reads above a certain baseline, counteracting drugs can be taken. 

27.5.1 DIFFERENCES BETWEEN HYPERBRANCHED STRUCTURES AND DENDRIMERS 

Dagani (1996) discussed and compared hyperbranched polymers and dendrimers. Hyperbranched 

polymers have highly branched structure, but unlike dendrimers, their structure is neither regular 

nor highly symmetrical. The dendrimers were obtained by careful stepwise growth of successive 

layers or generations. Hyperbranched polymers are obtained in a single step by polycondensation of 

A 2B monomers that contain two reactive groups of type A and one of type B. Functional groups A 

and B are selected in a way that they react with each other to form a covalent bond. Structure of 

hyperbranched polymers is thought to be intermediate between those of linear polymers and 

dendrimers. Because of steric constraints and statistical nature of coupling steps, hyperbranched 

polymers have irregular structures as some of the A groups remain unreacted, leading to incorporation 

of linear segments (i.e., monomers units coupled at two rather than three points) (Frechet 

1994). Voit, Turner, and Mourey (1993) showed that viscosity-MW behavior of the hyperbranched 

polymers does not show a maximum as observed in the case of dendrimers. Instead, they obey the 

Mark–Hownick–Sakunada relation. Although their viscosities are anomalously lower when 

compared to linear polymers of comparable size. Analogous to dendrimers, A 2B hyperbranched 

polymers had essentially same number of reactive groups as they have monomeric units. However, 

unlike dendrimers, these functional groups are not all located at chain ends (Frechet 1994). 

27.5.2 CLASSIFICATION OF DENDRIMERS 


The dendrimers that are most studied and reported up to higher generations (Bosman, Janssen, and 

Meijer 1999) can be classified chemically viz. peptide dendrimers that were patented and described 

upto higher generations by Denkewalter in the early 1980s. Those can well be characterized by 

size-exclusion chromatography only. PAMAM dendrimers are another class of dendritic structures 

that have been thoroughly investigated and had received wide spread attention. These were divergently 

constructed, e.g., PAMAM described by Tomalia and arborol systems of Newkome. 

Polypropyleneimine dendrimers were produced by Vogtle by divergent synthesis. It was also 

produced by heterogeneously catalyzed hydrogenation by De Brabander-van-den Berg and 

Meijer (1993) and is also studied widely. Frechet produced Poly-ether dendrimers by convergent 

procedures (Wooley et al. 1994). These also include Poly(aryl ether) dendrimers. Moore produced 

Phenyl acetylene dendrimers by convergent approach. Tetrathiofulvalene (TTF) 21-glycol 



dendrimers are another class of dendrimers that are intra-dendrimer aggregates of TTF cation 

radicals (Christensen et al. 1998). These are evident by spectrochemical studies of partially 

oxidized TTF units. These were prepared by convergent approaches. Pentaporphyrin is another 

starburst porphyrin polymer that is considered as a first generation dendrimers (Norsten and Branda 

1998). These had hybrid properties of porphyrin and dendrimers where ether linkages are found for 

iterations. Aryl ester monodispersed dendrimers based on 1, 3, 5-benzene tricarboxylic acid were 

initially described by Miller, Kwock, and Neenan (1992) and prepared by convergent synthesis. 

This was possibly used for MWs standards as polymeric rheological modifier, molecular inclusion 

hosts, etc. Two types of dendrimeric nucleic acids have been studied, viz. those that are covalently 

connected, and those that self-assemble via complementary base-pairing schemes (Kim and 

Zimmerman 1998). An example of a covalent DNA dendrimer was reported, wherein a secondgeneration 

phosphodiester-based dendrimer derived from pentaerythritol held nine pentathymidylate 

units on the end-groups and either a 15-mer or 26-mer at the core. 

Additionally, many other types of interesting, valuable, aesthetically pleasing dendritic systems 

have been developed such as Dendrophanes that were first described by Diederich and coworkers. 

They called these phenyl methane-based cyclophane. These were designed as globular proteins. 

Metallodendrimers include Ruthenium-terpyridine complexed dendrimers by Newkome and 

coworkers, dendritic iron (II) complexes by Chow and coworkers, zinc-porphyrin dendrimers by 

Diederich and coworkers, etc. and they contain metal complexed in dendrimer structure. Polyamino 

phosphine containing dendrimers are phosphorous containing dendrimers. Mesogen functionalized 

carbosilane dendrimer involves functionalization by 36 mesogenic units attached through C-5 

spacer to liquid crystalline dendrimer that form smectic-A phase in a temperature range of 117– 

1308C. Dendritic box was based on construction of a chiral shell of protected amino acids onto 

polypropyleneimine dendrimers with 64 amino end groups. These monodispersed dendritic 

containers of nanometric dimensions have physically entrapped or locked in guest molecules 

(Jansen et al. 1994; Jansen and Meijer 1995). Toyokuni et al. (1994) described an example of 

carbohydrate dendrimers for tumor-associated carbohydrate antigens. They linked it without the 

use of macromolecular carrier or an adjuvant, and they conjugated it with starburst PAMAM 

dendrimers to elicit antibody responses. 

Dendrimeric systems can also be classified on the basis of physical characteristics of the system 

such as simple dendrimers having simple monomeric units; liquid crystalline dendrimers having 

mesogenic liquid crystalline substances; chiral dendrimers having optical activities; and micellar 

dendrimers having unimolecular micellar structures or water-soluble hyperbranched dendrimers of 

polyphenylene, etc. Hybrid dendrimers are combinations of dendritic and linear polymers in hybrid 

block or graft copolymers forms, whereas amphiphilic dendrimers are a class of globular dendrimers 

having unsymmetrical, but highly controlled, distributions of chain end chemistry. These are 

oriented at interfaces forming interfacial liquid membranes for stabilizing aqueous organic emulsion. 

These may be oriented under influence of external stimulus, e.g., electric field. 

27.5.3 FEATURES OF DENDRIMERS AS NANODRUGS 

The various features of dendrimeric formulations in use include cost and ease of manufacture for 

cost effective (compared with small molecule drugs), and such formulations can be scaled-up for 

Current Good Manufacturing Practices (cGMP) guidelines manufacture. These are effective and 

active against a wide range of diseases and for their novel modes of action. Also at therapeutic 

doses, the dendrimers are safe and stable as solids, and they are thus available in a variety of 

Pharmaceutical formulations. The formulations are reproducible and defined as far as identity is 

concerned. They appear in most cases as white to off-white powder. Starpharma-like Pharmaceuticals 

(www.starpharma.com/framemaster.htm) are creating value from such dendrimer-based 

nanotechnology. Their core business strategies include developing high-value dendrimer nanodrugs 

to address unmet market needs, partnering with pharmaceutical companies to create new 


534 

opportunities and solutions to problems with the application of dendrimer nanotechnology, 

extending core skills and know-how through licensing and partnering with others, and investing 

in non-pharmaceutical applications of dendrimer technology. Rational drug design is being used by 

the biotechnology and pharmaceutical industry to design novel drugs based on known biological 

targets (e.g., receptors involved in the pathogenesis of disease). In this context, Starpharma’s 

dendrimer technologies offer a unique ability to design functionalized, polyvalent dendrimers as 

defined species for specific biological targets. 

The dendrimers are being eyed as drug-delivery agents, micelle mimics, and nanoscale building 

blocks. There are also added possibilities for grafting of active groups on surface of dendrimer by 

using cross-linkers. Also fatty acids, and polymers such as poly(ethyleneglycol) (PEG), etc. can be 

attached to dendrimers due to –COOH and –NH 2 terminal active groups present on its surface. PEG 

coatings, also called as PEGylation, could also have many possible numerous advantages so could 

be selected for coating (Zeng and Zimmerman 1997). Gitsov and Fréchet (1996) published some 

preliminary results on macromolecules that change shape when the polarity of the solvent changes. 

These macromolecules consist of four long PEG chains extending out from a central carbon atom. 

Each of these hydrophilic PEG arms is terminated with a hydrophobic wedge-like dendritic group 

based on 3,5-dihydroxybenzyl alcohol. The arms are long enough and flexible enough to allow this 

hybrid star to assume several very different conformations in solution. In tetrahydrofuran, the 

hybrid star forms a unimolecular micelle that has a hydrophilic core of tightly packed PEG arms 

surrounded by a loose hydrophobic shell of dendritic wedges. In chlorinated solvents such as 

chloroform where both building blocks are readily soluble, the PEG arms (and their dendritic 

extremities) extend outward, making the core more accessible. In polar or aqueous media such 

as methanol, the hydrophilic PEG arms loop around the hydrophobic dendritic wedges, pushing the 

wedges into the star’s core and leaving an outer PEG layer exposed to the outside world. The hybrid 

stars rearrange their two types of components in this way to minimize the overall free energy of 

the system. 

A novel hybrid star prepared in Fréchet’s lab at the University of California Berkeley changes 

its conformation when the polarity of the solvent medium changes. In THF (tetrahydrofuran), the 

hydrophilic PEG core is compact, whereas in chloroform, it is more extended and open. In 

methanol, the hydrophilic PEG arms envelope the hydrophobic dendritic wedges, leading to two 

possible conformations. In one, the dendritic wedges are tightly packed toward the middle of the 

micelle with the PEG arms forming loops in the surrounding medium. In the other arrangement, 

each dendritic moiety is individually wrapped in a PEG arm and somewhat extended into the 

medium. Freemantle (1999) also explored blossoming of dendrimers in a conference that explored 

design, synthesis, structure, and potential applications of highly branched macromolecules. 

27.6 APPLICATIONS OF PEGYLATED DENDRIMERS 


Dendrimers are synthetic, nanoscale structures (1–100 nm) types of polyvalent pharmaceuticals 

(www.starpharma.com/framemaster.htm) that can be tailored for many pharmaceutical applications. 

Specialized chemistry techniques allow for precise control of the physical and chemical 

properties of dendrimers. They are constructed in a series of controlled steps (see below) that 

increase the number of small branching molecules around a central core molecule. The final 

generation of molecules added to the growing structure makes up the polyvalent surface of the 

dendrimer. The core branching and surface molecules are chosen to give desired properties and 

functions. Dendrimers allow researchers, for the first time, to produce highly defined and biocompatible 

nanoscale objects built from the bottom up. This high definition enables unique 

functionality in life sciences applications. For readers interested in a historical account, terminology, 

nomenclature, or other background material, an excellent monograph by Newkome, 

Moorefield, and Vogtle (1996) is recommended. Zeng and Zimmerman (1997) broadly reviewed 



applications of dendrimers in bioorganic chemistry that have been reported in the past year, and, 

although somewhat different in focus, it updates earlier reviews published. An excellent review of 

dendrimers that act as biological mimics was also published by Smith and Diederich (1998). 

Beyond discussing a couple of results in the area of biomimetic dendrimers, this review highlighted 

various work on dendrimers whose monomer units are derived from carbohydrates, amino acids, or 

nucleotides and also several promising biomedical applications under active investigation. 

27.6.1 ENHANCED BIOCOMPATIBILITY FOR DRUG DELIVERY 

PAMAM dendrimers having poly(ethylene glycol) (PEG) grafts were designed as a novel drug 

carrier that possesses an interior for the encapsulation of drugs and a biocompatible surface. PEG 

monomethyl ether with the average MW of 550 or 2000 Da was combined to essentially every 

chain end of the dendrimer of the third or fourth generation via urethane bond (Kojima et al. 2000). 

The PEG-attached dendrimers encapsulating anti-cancer drugs adriamycin and methotrexate were 

prepared by extraction with chloroform from mixtures of the PEG-attached dendrimers and varying 

amounts of the drugs. Their ability to encapsulate these drugs increased with increasing dendrimer 

generation and chain length of PEG grafts. Among the PEG-attached dendrimers prepared, the 

highest ability was achieved by the dendrimer of the fourth generation having the PEG grafts with 

the average MW of 2,000 Da that could retain 6.5 adriamycin molecules or 26 methotrexate 

molecules/dendrimer molecule. The methotrexate-loaded PEG-attached dendrimers slowly 

released the drug in an aqueous solution of low ionic strength. However, in isotonic solutions, 

methotrexate and doxorubicin were readily released from the PEG-attached dendrimers. 

Padilla et al. (2002) described the use of high MW polymers (O 20,000 Da) as soluble drug 

carriers to improve drug targeting and therapeutic efficacy. They evaluated various dendritic architectures 

composed of a polyester dendritic scaffold based on the monomer unit 2,2bis(hydroxymethyl) 

propanoic acid for their suitability as drug carriers both in vitro and in vivo. 

These systems are both water-soluble and non-toxic. In addition, the potent anti-cancer drug 

doxorubicin was covalently bound via a hydrazone linkage to a high MW 3-arm poly(ethylene 

oxide) (PEO)-dendrimer hybrid. Drug release was a function of pH, and the release rate was more 

rapid at pH ! 6. The cytotoxicity of the DOX-polymer conjugate measured on multiple cancer 

lines in vitro was reduced but not eliminated, indicating that some active doxorubicin was released 

from the drug polymer conjugate under physiological conditions. Furthermore, biodistribution 

experiments show little accumulation of the DOX-polymer conjugate in vital organs, and the 

serum half-life of doxorubicin attached to an appropriate high MW polymer has been significantly 

increased when compared to the free drug. Therefore, this new macromolecular system exhibits 

promising characteristics for the development of new polymeric drug carriers. 

Kolhe et al. (2003) attached PEG grafts to the terminal amine groups of the dendrimer and 

enhanced biocompatibility of higher generation PAMAM dendrimer. Similarly, Bhadra et al. 

(2003) described the use of uncoated and PEGylated newer PAMAM dendrimers for delivery of 

an anti-cancer drug 5-fluorouracil. The 4.0 G PAMAM dendrimer was PEGylated using N-hydroxysuccinimide-activated 

carboxymethyl MPEG-5000. The PEGylation of the systems was found to 

have increased their drug-loading capacity, reduced their drug release rate, and hemolytic toxicity. 

TEM study revealed quite uniform surface of the systems. The systems were found suitable for 

prolonged delivery of an anti-cancer drug by in vitro and blood-level studies in albino rats without 

producing any significant hematological disturbances. PEG modification has been found to be 

suitable for modification of PAMAM dendrimers for reduction of drug leakage and hemolytic 

toxicity. This, in turn, could improve drug-loading capacity and stabilize such systems in the 

body. The study suggests use of such PEGylated dendrimeric systems as nanoparticulate depot 

type of system for drug administration (Figure 27.3). 

The entrapment of 5-fluorouracil in PEG modification dendrimers significantly increased by 12 

times because of more sealing of dendrimeric structure by PEG coating at the peripheral portions of 


536 

FIGURE 27.3 Chemical structural representation of PEG coated PAMAM loaded with 5-fluorouracil. 

(Adapted from Bhadra, et al., Int. J. Pharm., 257, 111, 2003. With permission.) 

dendrimers as coat that prevented drug release by enhancing complexation probably by steric and 

electronic effects of the additional functional groups made available by MPEG (Figure 27.4). This, 

in turn, also reduced the rate of drug release from such systems across dialysis membrane as 

observed by their release profile. The PAMAM dendrimers on PEGylation behaved similar to 

Cumulative % drug release 

120 

100 

80 

60 

40 

20 

0 

4.0 G non-PE Gylated 4.0G PEGylated 

1 2 3 4 5 6 7 24 48 72 96 120 144 

Time interval (hr.) 


FIGURE 27.4 Comparative cumulative percentage release of 5-fluorouracil from PEG coated and uncoated 

PAMAM dendrimers. (Adapted from Bhadra, et al., Int. J. Pharm., 257, 111, 2003. With permission.) 



(Concentration in blood (μg/ml) 

25 

20 

15 

10 

5 

0 

0 100 200 300 400 500 600 700 800 

Time (min) 

spherical unimolecular polymer micelles with chains of hydrated methoxy-poly(ethyleneglycol) on 

its surface as coat. This increased the entrapment capacity of dendrimers and made them act as 

nanometric containers carrying drugs. Such PEGylated dendrimers can be suggested for sustaining 

delivery of the drug 5-fluorouracil as also observed by the blood level data where blood level was 

much prolonged and was detectable up to 12 h. The formulations were found to be following 

sustained release characteristics for 5-FU (Figure 27.5) as shown by the relative increase in 

MRT for both non-PEGylated and PEGylated drug-dendrimer complexes as compared to plain 

drug solution (6.024 and 13.31 times, respectively). The release rate, however, was found to have 

increased in vivo as compared to in vitro data, possibly because of the metabolism by the enzymes 

and hydrolysis in the body. The blood level of the drug was found to be lower in case of PEGylated 

systems than that of non-PEGylated as a result of slower release rates of the drug similar to the trend 

found in vitro for the drug release. This might also be attributed to better cellular penetration and 

adhesion properties of the PEG chains of PEGylated systems that might lead to entanglement of 

such systems within capillary fenestration from which such systems might release drugs acting as 

nanoparticulate depot type carrier present in blood circulation. 

27.6.2 MICELLAR MEANS OF ENHANCING SOLUBILITY 

5-FU-PAMAM dend (non-PE Gylated) 

5-FU-PAMAM dend (PE Gylated) 

FIGURE 27.5 Blood level data of 5-fluorouracil from PEG coated and uncoated PAMAM dendrimers. 

(Adapted from Bhadra, et al., Int. J. Pharm., 257, 111, 2003. With permission.) 

Chapman and coworkers prepared amphiphilic copolymers derived from linear PEO and tBoCterminated 

poly-a,2-lysine dendrimers. The use of PEO as a platform for the dendrimers synthesis 

greatly facilitated separation because product up to fourth generation could be precipitated from 

reaction mixtures by ether. Surface tension method was employed to determine Critical Micelle 

Concentration (CMC) and micelle formation behavior. The surface of the aggregates was suggested 

to be highly compact. The existence of transition concentration for dye solubility of dye orange OT 

by G4 hydramphiphiles in water supports the micellar behavior (Chapman et al. 1994). 

Frechet and coworkers also synthesized another novel class of amphiphilic star copolymers. A 

four-armed PEG star was attached to four polyether dendrons derived from a pentaerythritol core. 

Results of SEC/VISC and 1 HNMR studies indicated the formation of unimolecular micelles in 

chloroform, tetrahydrofuran, or methanol, but with strikingly different structures. The star copolymers 

could self organizes into different micellar structures as a function of the environment. A 

potential application of these stimuli responsive copolymers involves their uses as solvent specific 

encapsulation agents (Gitsov and Frechet 1996). 


538 

Water-soluble dendritic unimolecular micelles as potential drug delivery agents had been 

explored by Liu et al. (2000) with the hydrophobic core surrounded by hydrophilic shell. These 

were prepared by coupling dendritic hypercores with PEG-mesylates. The monomeric core that was 

selected to build it was 4,4-bis(4 0 hydroxy phenyl) pentanol as this larger monomeric unit provides 

flexibility to the dendritic structures while contributing to the container capacity of the overall 

structure. Four generations of dendritic hypercores with 6, 12, 24, and 48 phenolic end groups were 

prepared. Subsequent coupling reaction with PEG-mesylates afforded four generations of dendritic 

unimolecular micelles (Gitsov and Fréchet 1993). The container property was demonstrated by 

pyrene solubilization in aqueous solution. In general, copolymers containing low generation 

dendrons, especially G1, tended to form unimolecular micelle, whereas G2 and G3 copolymers 

formed multi molecular micelles, presumably driven by hydrophobic effects and p–p interaction 

between dendritic blocks. Coexistence of two well-separated peaks in the SEC indicated a slow 

exchange process between these two species of different sizes. 

Ooya, Lee, and Park (2003) developed new methods and pharmaceutical compositions to 

increase the aqueous solubility of paclitaxel (PTX), a poorly water-soluble drug. Graft and starshaped 

graft polymers consisting of PEG400 graft chains increased the PTX solubility in water by 

three orders of magnitude. Polyglycerol dendrimers (dendriPGs) dissolved in water at high concentrations 

without significantly increasing the viscosity and at 80 wt% were found to increase the 

solubility of PTX 10,000-fold. The solubilized PTX was released from graft polymers, star-shaped 

graft polymers, and the dendriPGs into the surrounding aqueous solution. The release rate was a 

function of the star shape and the dendrimer generation. The availability of the new graft, star, and 

dendritic polymers having ethylene glycol units should permit development of novel delivery 

systems for other poorly water-soluble drugs. 

27.6.3 ENHANCING PHARMACOKINETIC PROPERTIES 


New polyester dendrimer-PEO bow-tie hybrids were evaluated for their potential as drug delivery 

vehicles and to explore the effect of MW and architecture on their pharmacokinetic properties 

(Elizabeth et al. 2005). In vitro experiments showed that these polymers were non-toxic to MDA- 

MB-231 cells and that they degraded to lower MWs at the normal physiological pH of 7.4 and at the 

mildly acidic pH of 5.0 that may be encountered upon uptake of these molecules by endocytosis and 

subsequent trafficking to endosomes and lysosomes. Biodistribution studies showed that all carriers 

with MWs of 40,000 and greater had long plasma circulation times, whereas those with lower MWs 

were cleared more rapidly with significant quantities excreted in the urine. In general, it was found 

that the more branched [G-3] polymers exhibited increased circulation times and decreased renal 

clearance relative to the less branched polymers, a property likely attributable to their decreased 

flexibility and resulting difficulty in passing through glomerular pores. It was also demonstrated that 

the more branched structures provide increased levels of steric protection for the payload on the 

drug-carrying dendron at the core of the molecule. High levels of tumor accumulation were found 

for two [G-3] polymers in mice bearing subcutaneous B16F10 melanoma. Overall, the features of 

these new branched carriers include degradability, lack of toxicity, long circulation half-lives, and 

high levels of tumor accumulation make them very promising polymers for therapeutic applications. 

High MW polymers have shown promise in terms of improving the properties and the 

efficacy of low MW therapeutics. However, new systems that are highly biocompatible are biodegradable, 

have well-defined MW, and have multiple functional groups for drug attachment are still 

needed. The biological evaluation of a library of eight polyester dendrimer-PEO bow-tie hybrids is 

described here. 

The group of evaluated polymers was designed to include a range of MWs (from 20,000 to 

160,000) and architectures with the number of PEO arms ranging from two to eight. In vitro 

experiments revealed that the polymers were non-toxic to cells and were degraded to lower MW 

species at pH 7.4 and pH 5.0. Biodistribution studies with 125 I-radiolabeled polymers showed that 



the high MW carriers (O40,000) exhibited long circulation half-lives. Comparison of the renal 

clearances for the four-arm versus eight-arm polymers indicated that the more branched polymers 

were excreted more slowly into the urine, a result attributed to their decreased flexibility. Because 

of their essentially linear architecture that does not provide for good isolation of the iodinated 

phenolic moieties, the polymers with two arms were rapidly taken up by the liver. The biodistributions 

of two long-circulating high MW polymers in mice bearing subcutaneous B16F10 tumors 

were evaluated, and high levels of tumor accumulation were observed. These new carriers are 

promising for applications in drug delivery and are also useful for improving the understanding of 

the effect of polymer architecture on pharmacokinetic properties (Elizabeth et al. 2005). 

27.6.4 DENDRIMER USE AS TRANSFECTION REAGENTS 

The recent completion of the human genome sequence has provided promising tools for defining 

cancer-specific targets, including genes and their expression patterns. In addition, the development 

in proteomics and achievements of high-throughput screenings offer great potential for discovering 

effective drugs, including DNA drugs, for gene therapy. Therefore, cancer therapy is no longer 

limited by the identification of genes; it is limited by an inability to deliver them efficiently to 

cancer cells, and the bottleneck in cancer gene therapy will soon be shifting from gene discovery 

to gene delivery (Luo 2004). The goal is to provide adequate amounts of drugs that are close to 

targeted cells because the actual destination for all gene delivery is the cell nucleus. Only drugs that 

are close to cells can be internalized. Once inside, intracellular barriers are considerably different. 

For example, diffusion is no longer the bottleneck. Rather, cytosol survival and nuclear targeting 

are more important (for non-viral gene delivery). An important and difficult task for cancer gene 

therapy is increasing the expression level of the therapeutic gene (that is usually toxic) in targeted 

tumor cells, but at the same time, avoiding its expression in normal tissue (Luo and Saltzman 2000). 

Methoxypoly(ethylene glycol)-block-poly(L-lysine) dendrimer was designed to form a watersoluble 

complex with plasmid DNA. The copolymer was synthesized by the liquid-phase peptide 

synthesis method (Choi et al. 1999). Agarose gel electrophoresis and DNase I protection assay 

proved that this linear polymer/dendrimer block copolymer spontaneously assembled with plasmid 

DNA, forming a water-soluble complex that increased the stability of the complexed DNA. Atomic 

force microscopy of the complex was evaluated at various charge ratios, showing that the copolymer/DNA 

complex was like a globular shape. 

Choi et al. (2000) synthesized and used a barbell-like triblock copolymer, poly(L-lysine) 

dendrimer-block-poly(ethylene glycol)-block-poly(L-lysine) dendrimer, for delivery of plasmid 

DNA by its self-assembly with the structures. A barbell-like ABA-type triblock copolymer, 

poly(L-lysine) dendrimer-block-poly(ethylene glycol)-block-poly(L-lysine) dendrimer 

(PLLD–PEG–PLLD), was synthesized by the liquid-phase peptide synthesis method. The selfassembling 

complex formation of the third and fourth generation of the copolymer with plasmid 

DNA was studied. 1 H NMR and matrix-assisted laser desorption/ionization-time-of-flight mass 

spectrometry (MALDI-TOF MS) were used for the characterization of the synthesized copolymer. 

The self-assembling behavior of the third and fourth generations of the copolymer with plasmid 

DNA was investigated by electrophoretic mobility shift assay, DNase I protection assay, and 

ethidium bromide exclusion assay. They observed great differences in the self-assembling ability 

of the third and fourth generations of the polymer. This suggests that the number of positively 

charged amines per polymer molecule should be an important factor for the potential for selfassembling 

complex formation with DNA. Atomic force microscopy (AFM) and z potentials were 

used for evaluating the shape, size distribution, and surface charge of the complexes at various 

charge ratios. From AFM images, it was observed that the shape of the complex was nearly 

spherical, and its size was about 50–150 nm in diameter. The in vitro cytotoxicity of the copolymer 

was compared with that of poly(L-lysine), poly(D-lysine), and polyethylenimine. 


540 

Luo et al. (2002) demonstrated a simple and successful synthetic approach to devise a highly 

efficient DNA delivery system with low cytotoxicity and low cost. PAMAM dendrimer is a highly 

efficient DNA delivery agent when compared to other chemical transfection reagents. Partially 

degraded, high-generation dendrimers offer even higher efficiency, presumably because of 

enhanced flexibility of the otherwise rigid dendrimer chains. It was hypothesized that chemical 

modification of low generation dendrimer with biocompatible PEG chains would create a conjugate 

of PAMAM core with flexible PEG chains that mimics the fractured high-generation dendrimer and 

produces high transfection efficiency. Generation 5 PAMAM dendrimer was modified with 

3400 MW PEG. The novel conjugate produced a 20-fold increase in transfection efficiency 

compared with partially degraded dendrimer controls. The cytotoxicity of PEGylated dendrimers 

was very low. This extremely efficient, highly biocompatible, low-cost DNA delivery system can 

be readily used in basic research laboratories and may find future clinical applications. 

This system can be readily used in basic research laboratories and may be modified and applied 

in future clinical usage. Furthermore, similar approaches could be adopted with other DNA delivery 

agents that may create new opportunities for using DNA as a drug via a synthetic delivery route. 

Mannisto et al. (2002) investigated the influence of shape, MW, and PEGylation of linear, 

grafted, dendritic, and branched poly(L-lysine) on their DNA delivery properties. DNA binding, 

condensation, complex size and morphology, cell uptake, and transfection efficiency were 

determined. Most poly(L-lysine) condense DNA linear polymers and are at the same time more 

efficient than most dendrimers. At low MWs of PLL, DNA binding and condensation were less 

efficient, particularly with dendrimers. PEGylation did not decrease DNA condensation of PLLs at 

less than 60% (fraction of MW) of PEG. PEGylation sterically stabilized the complexes but did not 

protect them from interaction with polyanionic chondroitin sulfate. Cell uptake of poly(L-lysine)/ 

DNA complexes was high and PEGylation increased the transfection efficacy. However, overall 

transfection level of poly(L-lysine) is low possibly as a result of the inadequate escape of the 

complexes from endosomes or poor release of DNA from the complexes. Physicochemical and 

biological structure-property relationships of poly(L-lysine) were demonstrated, but no clear correlations 

between the tested physicochemical determinants (size of complexes, zeta-potentials, 

condensation of DNA, and the shape of complexes) and biological activities were seen. Intracellular 

factors and/or still unknown features of DNA complexation may ultimately determine 

transfection activity with the carriers. 

Kim et al. (2004) synthesized and applied a novel triblock copolymer, PAMAM-block-PEGblock-PAMAM, 

as a gene carrier. PAMAM dendrimer is proven to be an efficient gene carrier 

itself, but it is associated with certain problems such as low water solubility and considerable 

cytotoxicity. Therefore, they introduced PEG to engineer a non-toxic and highly transfectionefficient 

polymeric gene carrier because PEG is known to convey water-solubility and biocompatibility 

to the conjugated copolymer. This copolymer could achieve self-assembly with plasmid 

DNA, forming compact nanosized particles with a narrow size distribution. Fulfilling expectations, 

the copolymer was found to form highly water-soluble polyplexes with plasmid DNA, showed little 

cytotoxicity despite its poor degradability, and finally achieved high transfection efficiency comparable 

to PEI. Consequently, they showed that an approach involving the introduction of PEG to 

create a tree-like cationic copolymer possesses a great potential for use in gene delivery systems. 

Logically, this approach could be extended to cancer treatment. 

27.6.5 STIMULI RESPONSIVE CARRIERS 


Gillies and Frechet (2002) designed and prepared a new polyester dendrimer, PEO hybrid systems, 

for drug delivery and related therapeutic applications. These systems consist of two covalently 

attached polyester dendrons where one dendron provides multiple functional handles for the attachment 

of therapeutically active moieties, whereas the other is used for attachment of solubilizing 

PEO chains. By varying the generation of the dendrons and the mass of the PEO chains, the MW, 



architecture, and drug loading can be readily controlled. The bow-tie shaped dendritic scaffold was 

synthesized using both convergent and divergent methods with orthogonal protecting groups on the 

periphery of the two dendrons. PEO was then attached to the periphery of one dendron using an 

efficient coupling procedure. A small library of eight carriers with MWs ranging from about 20 to 

160 kDa were prepared and characterized by various techniques, confirming their welldefined 

structures. 

Gillies, Jonsson, and Frechet (2004) used a pH-responsive micelle system, linear-dendritic 

block copolymers comprising PEO and either a polylysine or polyester dendron. These were 

prepared, and hydrophobic groups were attached to the dendrimer periphery by highly acid-sensitive 

cyclic acetals. These copolymers were designed to form stable micelles in aqueous solution at 

neutral pH and disintegrate into unimers at mildly acidic pH following loss of the hydrophobic 

groups upon acetal hydrolysis. Micelle formation was demonstrated by encapsulation of the fluorescent 

probe Nile Red, and the micelle sizes were determined by dynamic light scattering. The 

structure of the dendrimer block, its generation, and the synthetic method for linking the acetal 

groups to its periphery all had an influence on the CMC and the micelle size. The rate of hydrolysis 

of the acetals at the micelle core was measured for each system at pH 7.4 and pH 5, and it was found 

that all systems were stable at neutral pH but that they underwent significant hydrolysis at pH 5 over 

several hours. The rate of hydrolysis at pH 5 was dependent on the structure of the copolymer, most 

notably the hydrophobicity of the core-forming block. To demonstrate the potential of these 

systems for controlled release, the release of Nile Red as a model payload was examined. At pH 

7.4, the fluorescence of micelle-encapsulated Nile Red was relatively constant, indicating it was 

retained in the micelle, and at pH 5, the fluorescence decreased, consistent with its release into the 

aqueous environment. The rate of release was strongly correlated with the rate of acetal hydrolysis 

and was controlled by the chemical structure of the copolymer. The mechanism of Nile Red release 

was investigated by monitoring the change in size of the micelles over time at acidic pH. Dynamic 

light scattering measurement showed a size decrease over time, eventually reaching the size of a 

unimer, providing evidence for the proposed micelle disintegration. 

27.6.6 ALTERATION OF TOXICITY 

In one of the studies Bhadra et al. (2003) found that PEGylation significantly decreased the 

hemolysis of red blood cells (RBCs) to below 5% in case of whole generation amine terminated 

charged dendrimers having about 15.3–17.3% hemolytic toxicity, similar to negligible in case of 

half generations of carboxylic acid terminated dendrimers. This was due to inhibition of interaction 

of RBCs with the charged quaternary ammonium ion as determined by interaction with RBCs. The 

hemolytic toxicity of the dendrimers was enough to preclude its use as a drug delivery system. The 

toxicity was due to the poly-cationic nature of the PAMAM dendrimers that was also responsible 

for their cytotoxicity (Malik et al. 2000), particularly in the case of whole generation amine 

terminated charged dendrimers, but not in the case of half generations of carboxylic acid 

terminated dendrimers. 

The PEGylation of the dendrimers were found to have also considerably lower rate of hemolysis 

of the RBCs because of significantly reduced interaction of RBCs with the charged quaternary 

ammonium ion that is generally present on the amine terminated whole generations of dendrimers. 

Also, the hematological study was undertaken to assess the relative effects of the PEGylated and 

non-PEGylated systems as compared to the plain drug on various blood parameters. The blood 

parameters undergoing major changes as to normal values of blood levels are RBC count, WBC 

count, and differential lymphocytes count. The RBC count of non-PEGylated 5-FU-dendrimers was 

found to have decreased below normal values by about 1!10 6 /ml RBCs to 2!10 6 /ml RBCs as 

against similar PEGylated systems. The WBC count of non-PEGylated 5-FU-dendrimer complex 

increased by 3!10 3 /ml cells to 4!10 3 /ml cells as compared to normal values. However, for 5 

FU-PEGylated dendrimer complexes, the increase was by 1!10 3 /ml cells to 1.5!10 3 /ml WBCs 


542 


as compared to normal count in controlled group. The increase in WBC count is significant in case 

of non-PEGylated systems. Similarly, relatively greater increase in lymphocyte count was observed 

by non-PEGylated dendrimer-drug complexes that were by about 2!10 3 /ml cells. This was similar 

to blood toxicity and cytotoxicity effects of acrylates nanoparticulates (Malik et al. 2000), which are 

known to be stimulating the macrophage level and WBC count. This was found to be true in the 

case of these non-PEGylated drug-dendrimeric systems but not in case of PEGylated systems, 

conforming the trends of the PEGylated carriers (Veronese et al. 1989; Papahadjopoulus et al. 

1991; Phillips et al. 1999) undergoing lesser phagocytic uptake. 

Neerman et al. (2004) also showed that dendrimers based on melamine could reduce the organ 

toxicity of solubilized cancer drugs administered by intraperitoneal injection. Methotrexate and 

6-mercaptopurine, both FDA approved anti-cancer drugs, are known hepatotoxins. The solubility 

of these molecules can be increased by mixing them with a dendrimer based on melamine. C3H 

mice were administered subchronic doses of methotrexate or 6-mercaptopurine with and without a 

solubilizing dendrimer. Forty-eight hours after dosing, the mice were sacrificed, and serum was 

collected for biochemical analyses. The levels of alanine transaminase, ALT, were used to probe 

liver damage. When the drugs are encapsulated by the dendrimer, a significant reduction in hepatotoxicity 

is observed; ALT levels from the rescued groups (drug C dendrimer) were 27% 

(methotrexate) and 36% (6-mercaptopurine) lower than those of animals treated with the 

drug alone. 

Initial studies of a generation 3, cationic dendrimer bearing 24 primary amino groups revealed 

that this dendrimer was toxic both in vitro and in vivo. Extensive liver damage was observed when 

mice were challenged with subchronic doses of this dendrimer. Other groups hypothesized that 

surface group charge was the determining factor in predicting a dendrimer’s toxicity as cationic 

surface groups will electrostatically adhere to cell surface, inducing bulk adhesion that leads to the 

cell’s demise. Surface group modification of a generation 3, cationic dendrimer bearing 48 primary 

amino groups to possess anionic or neutral surface groups had a significant impact on both in vitro and 

in vivo toxicity. The introduction of neutral PEG groups afforded the most protection in vitro and 

in vivo. Shielding of the dendrimer’s charged groups attenuates the bulk adhesion of these molecules, 

leading to a decrease in the toxicity of these dendrimers. It appeared that PEGylation afforded the 

most protection against toxicity of dendrimers based on melamine. As a result, the candidate vehicle 

was subjected to a series of analyses. Albumin interaction studies showed that this dendrimer 

interacted with albumin that could have a profound effect on the dendrimer’s pharmacokinetic 

parameters. In vitro drug release profiles of the anti-cancer drugs methotrexate and paclitaxel 

revealed that methotrexate was rapidly released while paclitaxel showed a more controlled, sustained 

release. This discrepancy in drug release was attributed to disparities in the drug’s hydrophobicity; 

the more hydrophobic drug, paclitaxel, formed a stronger association with the dendrimer’s hydrophobic 

core. As a result, release was more controlled. In terms of the in vitro cytotoxicity of these 

dendrimer/drug formulations, free drug was more toxic than the dendrimer-associated drug. This was 

attributed to differences in the cellular uptake offree versus dendrimer-associated drug. To assess if a 

difference in cellular uptake existed between free and dendrimer associated substrates, cells exposed 

to free rhodamine B showed higher intracellular levels of dye than those cells exposed to dendrimerassociated 

rhodamine B. The preliminary in vivo biodistribution screen of the Cy5.5-labeled 

dendrimer showed that the dendrimer was still present in the body 24 h post injection while exhibiting 

a high degree of liver accumulation after such time (https://txspace.tamu.edu/dev-xml/bitstream/ 

1969.1/5330/1/etd-tamu-2005A-TOXI-Neerman.pdf). 

Chen et al. (2004) designed a small library of dendrimers prepared from a common precursor 

that is available in 5 g scale in five linear steps at 56% overall yield. The precursor is a generation 

three dendrimer that displays 48 peripheral sites by incorporating AB4 surface groups. Manipulation 

of these sites provided six dendrimers that vary in the chemistry of the surface group (amine, 

guanidine, carboxylate, sulfonate, phosphonate, and PEGylated) that were evaluated for hemolytic 

potential and cytotoxicity. Cationic dendrimers were found to be more cytotoxic and hemolytic 



than anionic or PEGylated dendrimers. The PEGylated dendrimer was evaluated for acute toxicity 

in vivo. No toxicity, neither mortality nor abnormal blood chemistry, based on blood urea nitrogen 

levels or alanine transaminase activity was observed in doses up to 2.56 g/kg i.p. and 

1.28 g/kg intavenously. 

27.6.7 EFFECTIVE PHOTODYNAMIC THERAPY 

Photosensitizers play a crucial role in the photodynamic therapy (PDT) of cancer. Zhang et al. 

(2003) observed that the use of PIC micelles as a delivery system reduced the dark toxicity of the 

cationic dendrimer porphyrin, probably because of the biocompatible PEG shell of the micelles. 

They compared relatively low cellular uptake of dendrimer porphyrin [NH2CH2CH2NHCO]32- 

DPZn incorporated in the PIC micelle as was observed, yet the latter exhibited enhanced 

photodynamic efficacy on the Lewis Lung Carcinoma (LLC) cell line. In this study, a third-generation 

aryl ether dendrimer porphyrin with 32 primary amine groups on the periphery, 

[NH2CH2CH2NHCO]32DPZn, and pH-sensitive, polyion complex micelles (PIC) composed of 

the porphyrin dendrimer and PEG-b-poly(aspartic acid) were evaluated as new photosensitizers 

(PSs) for PDT in the LLC cell line. The preliminary photophysical characteristics of [NH2CH2- 

CH2NHCO]32DPZn and the corresponding micelles were investigated. Electrostatic assembly 

resulted in a red-shift of the Soret peak of the porphyrin core and the enhanced fluorescence. 

Jang et al. (2005) described a supramolecular nanocarrier of anionic dendrimer porphyrins 

with cationic block copolymers modified with PEG to enhance intracellular photodynamic efficacy. 

27.6.8 DENDRITIC GELS 

Luman, Smeds, and Grinstaff (2003) developed a process of high-yield convergent synthesis of 

dendrons, dendrimers, and dendritic-linear hybrid macromolecules composed of succinic acid, 

glycerol, and PEG. This convergent synthesis relies on two orthogonal protecting groups, 

namely, the benzylidene acetal (bzld) for the protection of the 1,3-hydroxyls of glycerol and the 

tert-butyldiphenylsialyl (TBDPS) ester for protection of the carboxylic acid of succinic acid. These 

novel polyester dendritic macromolecules are entirely composed of building blocks known to be 

biocompatible or degradable in vivo to give natural metabolites. Derivatization of the dendritic 

periphery with a methacrylate affords a polymer that can be subsequently photo-cross-linked. The 

three-dimensional cross-linked gels formed by ultraviolet irradiation are optically transparent with 

mechanical properties dependent on the initial cross-linkable dendritic macromolecule. 

27.6.9 LIGAND-BASED DRUG DELIVERY 

Architectural features of synthetic ligands were systematically varied to optimize inhibition of mast cell 

degranulation initiated by multivalent crossing of IgE-receptor complexes Baird et al. (2003). A series 

of ligands were generated by end-capping PEG polymers and amine-based dendrimers with the hapten 

2,4-dinitrophenyl (DNP). These were used to explore the influence of polymeric backbone length, 

valency, and hapten presentation on binding to anti-DNP IgE and inhibition of stimulated activation of 

RBL cells. Monovalent MPEG(5000)-DNP (IC(50)Z50 nM), bivalent DNP-PEG(3350)-DNP 

(IC(50)Z8 nM), bismonovalent MPEG(5000)-DNP(2) (IC(50)Z20 nM), bisbivalent DNP(2)- 

PEG(3350)-DNP(2) (IC(50)Z3 nM), and DNP(4)-dendrimer ligands (IC(50)Z50 nM) all effectively 

inhibit cellular activation caused by multivalent antigen, DNP-bovine serum albumin. For 

different DNP ligands, evidence is available for more effective inhibition because of preferential 

formation of intra-IgE cross-links by bivalent ligands of sufficient length, self-association of monovalent 

ligands with longer tails, and higher probability of binding for bisvalent ligands. They also 

showed that larger DNP(16)-dendrimers of higher valency trigger degranulation by cross-linking 


544 

IgE-receptor complexes, whereas smaller DNP-dendrimers are inhibitory. Therefore, features of 

synthetic ligands can be manipulated to control receptor occupation, aggregation, and inhibition of 

the cellular response. 

27.6.10 INCREASINGLY EFFECTIVE BORON NEUTRON CAPTURE THERAPY 

A recent methodology in cancer treatment is the boron neutron-capture therapy (BNCT) 

(Hawthorne 1993). In this therapy, the generation of cytotoxic and energetic products from 

nuclear fission reactions of low-energy neutrons and 10 B nuclei is used to destroy malignant 

cells. An efficient agent for the therapy is water-soluble and has a high local density of boron 

clusters, requirements that are met for several synthesized boron-containing and water-soluble 

dendrimers. Qualmann et al. (1996a, 1996b) have, in addition, introduced antigen selectivity by 

coupling a lysine-based boronated dendrimer to antibody fragments (Fab¢). In these agents, the 

attachment of a poly(ethylene glycol)(PEG) moiety is necessary to keep the conjugates watersoluble. 

The covalent nature of the boronated Fab¢ fragments leads to a better stability of these 

conjugates as compared to, for example, borate-coated polystyrene beads. 

Successful treatment of cancer by boron neutron capture therapy (BNCT) requires the selective 

delivery of 10 B to constituent cells within a tumor. The expression of the folate receptor is amplified 

in a variety of human tumors and potentially might serve as a molecular target for BNCT. Shukla et 

al. (2003) investigated the possibility of targeting the folate receptor on cancer cells using folic acid 

conjugates of boronated poly(ethylene glycol) (PEG) containing third generation PAMAM dendrimers 

to obtain 10 B concentrations necessary for BNCT by reducing the uptake of these conjugates 

by the reticuloendothelial system. First, they covalently attached 12–15 decaborate clusters to third 

generation PAMAM dendrimers. Varying quantities of PEG units with varying chain lengths were 

then linked to these boronated dendrimers to reduce hepatic uptake. Among all prepared combinations, 

boronated dendrimers with 1–1.5 PEG(2000) units exhibited the lowest hepatic uptake in 

C57BL/6 mice (7.2–7.7% injected dose (ID)/g liver). Therefore, two folate receptor-targeted boronated 

third generation PAMAM dendrimers were prepared: one containing approximately 15 

decaborate clusters and approximately 1 PEG(2000) unit with folic acid attached to the distal 

end, and the other containing approximately 13 decaborate clusters with, approximately one 

PEG(800) unit with folic acid attached to the distal end. In vitro studies using folate receptor 

(C) KB cells demonstrated receptor-dependent uptake of the latter conjugate. Biodistribution 

studies with this conjugate in C57BL/6 mice bearing folate receptor (C) murine 24JK-FBP 

sarcomas resulted in selective tumor uptake (6.0% ID/g tumor), but also high hepatic (38.8% 

ID/g) and renal (62.8% ID/g) uptake, indicating that attachment of a second PEG unit and/or 

folic acid may adversely affect the pharmacodynamics of this conjugate. 

27.6.11 IMPROVED MRI CONTRAST AGENTS 


Macromolecules conjugated with polyethylene glycol (PEG) acquire more hydrophilicity, resulting 

in a longer half-life in circulation and lower immunogenicity. Kobayashi et al. (2001) synthesized 

two novel conjugates for MRI contrast agents from a generation-4 polyamidoamine dendrimer 

(G4D), 2-(p-isothiocyanatobenzyl)-6-methyl-diethylenetriaminepentaacetic acid (1B4M), and one 

or two PEG molecules with a MW of 20,000 Da (PEG(2)-G4D-(1B4M-Gd)(62) (MW: 96 kD), 

PEG(1)-G4D-(1B4M-Gd)(63) (MW: 77 kD). Their pharmacokinetics, excretion, and properties as 

vascular MRI contrast agents were evaluated and compared with those of G4D-(1B4M-Gd)(64) 

(MW: 57 kD). PEG(2)-G4D-(1B4M-Gd)(62) remained in the blood significantly longer and accumulated 

significantly less in the liver and kidney than the other two preparations (P!0.01). 

Although the blood clearance was slower, PEG(2)-G4D-(1B4M-Gd)(62) was excreted more 

readily without renal retention than the other two preparations. The positive effects of PEG 



conjugation on a macromolecular MRI contrast agent were found to be prolonging retention in the 

circulation, increased excretion, and decreased accumulation in the organs. 


Nanotechnology refers to the interactions of cellular and molecular components and engineered 

materials—typically clusters of atoms, molecules, and molecular fragments—at the most elemental 

level of biology. Such nanoscale objects—typically, though not exclusively, with dimensions 

smaller than 100 nm—can be useful by themselves or as part of larger devices, containing multiple 

nanoscale objects. At the nanoscale, the physical, chemical, and biological properties of materials 

fundamentally differ, and often nanotechnology can change the very foundations of cancer diagnosis, 

treatment, and prevention. The novel nanodevices are capable of many clinically important 

functions, including detecting cancer at its earliest stages, pinpointing its location within the body, 

delivering anti-cancer drugs specifically to malignant cells, and in killing malignant cells. As these 

nanodevices are evaluated in clinical trials, researchers envision that nanotechnology will serve as 

multifunctional tools that will not only be used with any number of diagnostic and therapeutic 

agents, but will also change the very foundations of cancer diagnosis, treatment, and prevention 

(NCI 2004). 

Such nanoparticulate carriers also include PEG conjugated dendrimers. These are globular, 

hyperbranched polymers possessing a high concentration of surface functional groups and internal 

cavities. These unique features make them very useful in many biomedical applications, especially 

as carrier molecules. There are many such technological opportunities associated with the dendrimers. 

Starpharma-like companies used these dendrimers to build large active compounds that 

present a polyvalent array to receptors, and others platform opportunities for the delivery of 

small molecules as a toolbox for rational drug design and other nanotechnology applications in 

life sciences. There are many product opportunities in dendrimers and their use such as with 

VivaGele, a topical microbicide gel for prevention of HIV and other STDs in women. They can 

act as novel chemotherapeutic agents, angiogenesis inhibitors, and they can be used for targeting a 

range of tropical and exotic diseases and as bio-defense agents in addition to their use for targeting a 

broad range of viral, respiratory, and cancer-like diseases in the future. Development of drug 

delivery systems with low toxicity and low cost is the key to a nanaotechnological approach to 

cancer treatment, including DNA-based concepts. Targeting the drugs to specific receptors 

exclusively expressed by tumor cells also appears to be an approach worth exploring. This biochemical 

approach is expected to pay rich dividends in the chemotherapy of cancer. Cytosol 

survival of anti-cancer agents and folate receptor-based approaches also hold promise. Carbon 

nanotubes offer yet another attractive possibility. 

REFERENCES 

Abou-Jawde, R. et al., An overview of targeted treatments in cancer, Clin. Ther., 25, 2121, 2003. 

Baird, E. J. et al., Highly effective poly(ethylene glycol) architectures for specific inhibition of immune 

receptor activation, Biochemistry, 42, 12739, 2003. 

Barar, F. S. K., Essential of Pharmacotherapeutics, S. Chand and Company Ltd, New Delhi, 2000. 

Bhadra, D. et al., Pegnology: A review of PEG-ylated systems, Pharmazie, 57, 5, 2002. 

Bhadra, D. et al., A PEGylated dendritic nanoparticulate carrier of fluorouracil, Int. J. Pharm., 257, 111, 2003. 

Bosman, A. W., Janssen, H. M., and Meijer, E. W., About dendrimers: Structure, physical properties and 

applications, Chem. Rev., 99, 1665, 1999. 

Brannon-Peppas, L. and Blanchette, J. O., Nanoparticle and targeted systems for cancer therapy, Adv. Drug 

Deliv. Rev., 56, 1649, 2004. 

Brigger, I., Dubernet, C., and Couvreur, P., Nanoparticles in cancer therapy and diagnosis, Adv. Drug Deliv. 

Rev., 54, 631, 2002. 


546 


Buckmann, A. F. and Morr, M., Functionalization of polyethylene glycol and mono-methoxy-polyethylene 

glycol, Makromol. Chem., 182, 1379, 1981. 

CCO Formulary, Liposomal Doxorubicin, Oct. 2003. 

Chapman, T. M. et al., Hydramphiphiles: Novel, linear, dendritic block copolymeric surfactants, J. Am. Chem. 

Soc., 116, 11195, 1994. 

Chari, R. V. J., Targeted delivery of chemotherapeutics: Tumor activated prodrug therapy, Adv. Drug Deliv. 

Rev., 31, 89, 1998. 

Chen, H. T. et al., Cytotoxicity, hemolysis, and acute in vivo toxicity of dendrimers based on melamine, 

candidate vehicles for drug delivery, J. Am. Chem. Soc., 126, 10044, 2004. 

Choi, Y. H. et al., Polyethylene glycol grafted poly-L-lysine as polymeric gene carrier, J. Controlled Release, 

54, 39, 1998. 

Choi, J. S. et al., Poly(ethylene glycol)-block-poly(L-lysine) dendrimer: Novel linear polymer/dendrimer block 

copolymer forming a spherical water-soluble polyionic complex with DNA, Bioconjugate Chem., 10, 

62, 1999. 

Choi, J. S. et al., Synthesis of a barbell-like triblock copolymer, poly(L-lysine) dendrimer-block-Poly(ethylene 

glycol)-block-Poly(L-lysine) dendrimer, and its self-assembly with plasmid DNA, J. Am. Chem. Soc., 

122, 474, 2000. 

Christensen, C. A. et al., Synthesis and electrochemistry of a tetrathiafulvalene (TTF)21-glycol dendrimer: 

Intradendrimer aggregation of TTF cation radicals, Chem. Commun., 509, 1998. 

Cui, Z., Hsu, C. H., and Mumper, R. J., Physical characterization and macrophage cell uptake of mannancoated 

nanoparticles, Drug Dev. Ind. Pharm., 29, 689, 2003. 

Dagani, R., Chemists explore potential of dendritic macromolecules as functional materials, CandEN 

Washington, Chem. and Eng. News, June 3, 1996. 

De Brabander-Van den Berg, E. M. M. and Meijer, E. W., Poly (propyleneimine) dendrimers: Large-scale 

synthesis by heterogeneously catalyzed hydrogenations, Angew. Chem. Int. Ed. Engl., 32, 1308, 1993. 

Dvorak, H. F. et al., Identification and characterization of the blood vessels of solid tumors that are leaky to 

circulating macromolecules, Am. J. Pathol., 133, 95, 1988. 

Elamanchili, P. et al., Characterization of poly(D,L lactic-co-glycolic acid) based nanoparticulate system for 

enhanced delivery of antigens to dendritic cells, Vaccine, 22, 2406, 2004. 

Elizabeth, R. et al., Biological evaluation of polyester dendrimer: Poly(ethylene oxide) “bow-tie” hybrids with 

tunable molecular weight and architecture, Mol. Pharmacol., 2, 129, 2005. 

Feng, S. S. and Chien, S., Chemotherapeutic engineering: Application and further development of chemical 

engineering principles for chemotherapy of cancer and other diseases, Chem. Eng. Sci., 58, 4087, 

2003. 

Frechet, J. M., Functional polymers and dendrimers: Reactivity, molecular architecture, and interfacial energy, 

Science, 263, 1710, 1994. 

Freemantle, M., Blossoming of dendrimers, Chem. & Engg. News London, Science/Technology, 77, 27, 1999. 

Freitas, R. A. Jr., Nanomedicine, In Basic Capabilities, Landes Bioscience, Georgetown, TX, I, (http://www. 

nanomedicine.com/NMI.htm), 1999. 

Freitas, R. A., Current status of nanomedicine and medical nanorobotics, J. Comput. Theor. Nanosci., 

2, 1, 2005. 

Gabizon, A. and Martin, F., Polyethylene glycol-coated (PEGylated) liposomal doxorubicin. Rationale for use 

in solid tumours, Drugs, 54, 15, 1997. 

Gillies, E. R. and Frechet, J. M., Designing macromolecules for therapeutic applications: Polyester dendrimerpoly(ethylene 

oxide) “bow-tie” hybrids with tunable molecular weight and architecture, J. Am. Chem. 

Soc., 124, 14137, 2002. 

Gillies, E. R., Jonsson, T. B., and Frechet, J. M., Stimuli-responsive supramolecular assemblies of lineardendritic 

copolymers, J. Am. Chem. Soc., 126, 11936, 2004. 

Gitsov, I. and Fréchet, J. M. J., Solution and solid state properties of hybrid linear-dendritic block copolymers, 

Macromolecules, 26, 6536, 1993. 

Gitsov, I. and Fréchet, J. M. J., Stimuli responsive hybrid macromolecules: Novel amphiphilic star copolymers 

with dendritic groups at the periphery, J. Am. Chem. Soc., 118, 3785, 1996. 

Grossfeld, G. D., Carrol, P. R., and Lindeman, N., Thrombospondin-1 expression in patients with pathologic 

state T3 prostate cancer undergoing radical prostatectomy: Association with p53 alterations, tumor 

angiogenesis and tumor progression, Urology, 59, 97, 2002. 



Harris, J. M. et al., Synthesis and characterisation of polyethylene glycol derivatives, J. Polymer Sci. Polymer 

Chem., 22, 341, 1984. 

Harris, J. M. and Chess, R. B., Effect of pegylation on pharmaceuticals, Nat. Rev. Drug Discov., 2, 214, 2003. 

Hawthorne, M. F., The role of chemistry in the development of boron neutron capture therapy of cancer, 


http://nano.med.umich.edu/Dendrimers.html. 

http://press2.nci.nih.gov/sciencebehind/nanotech. 

http://www.cancer.gov/clinicaltrials/results/dose-dense0604. 

Ishida, O. et al., Size dependent extravagation and interstitial localization of polyethylene glycol liposomes in 

solid tumour bearing mice, Int. J. Pharm., 190, 49, 1999. 

Jang, W. D. et al., Supramolecular nanocarrier of anionic dendrimer porphyrins with cationic block copolymers 

modified with polyethylene glycol to enhance intracellular photodynamic efficacy, Angew. 

Chem. Int. Ed. Engl., 44, 419, 2005. 

Jansen, J. F. G. A. and Meijer, E. W., The dendritic box: Shape-selective liberation of encapsulated guests, 

J. Am. Chem. Soc., 117, 4417, 1995. 

Jansen, J. F. G. A., Brabander Vanden Berg, E. M. M., and Meijer, E. W., Science, 266, 1226, 1994. 

Katre, N. V., The conjugation of protein with poly ethylene glycol and other polymers altering properties of 

proteins to enhance their therapeutic potential, Adv. Drug Deliv. Rev., 10, 91, 1993. 

Kim, A. et al., Pharmacodynamics of insulin in polyethylene glycol coated liposomes, Int. J. Pharmacol., 180, 

75, 1999. 

Kim, T. I. et al., PAMAM–PEG–PAMAM: Novel triblock copolymer as a biocompatible and efficient gene 

delivery carrier, Biomacromolecules, 5, 2487, 2004. 

Kim, Y. and Zimmerman, S. C., Applications of dendrimers in bio-organic chemistry, Curr. Opin. Chem. Biol., 

2, 733, 1998. http://biomednet.com/elecref/1367593100200733. 

Kobayashi, H. et al., Positive effects of polyethylene glycol conjugation to generation-4 polyamidoamine 

dendrimers as macromolecular MR contrast agents, Magn. Reson. Med., 46(4), 781, 2001. 

Kojima, C. et al., Synthesis of polyamidoamine dendrimers having poly(ethylene glycol) grafts and their 

ability to encapsulate anticancer drugs, Bioconjugate Chem., 11(6), 910, 2000. 

Kolhe, P. et al., Drug complexation, in vitro release and cellular entry of dendrimers and hyperbranched 

polymers, Int. J. Pharm., 259, 143, 2003. 

Kreuter, J., Nanoparticles, In Colloidal Drug Delivery Systems, Kreuter, J., Ed., Marcel Dekker, Inc., 

New York, 1994. 

Kukowska-Latallo, J. F. et al., Intravascular and endobronchial DNA delivery to murine lung tissue using a 

novel, nonviral vector, Hum. Gene Ther., 11, 1385, 2000. 

Langer, R., Biomaterials in drug delivery and tissue engineering: One laboratory’s experience, Acc. Chem. 

Res., 33, 94, 2000. 

Landers, J. J. et al., Prevention of influenza pneumonitis by sialic acid-conjugated dendritic polymers, J. Infect. 

Dis., 186, 1222, 2002. 

Liu, M., Kono, K., and Frechet, J. M., Water-soluble dendritic unimolecular micelles: their potential as drug 

delivery agents, J. Control. Rel., 65, 121, 2000. 

Luman, N. R., Smeds, K. A., and Grinstaff, M. W., The convergent synthesis of poly(glycerol-succinic acid) 

dendritic macromolecules, Chemistry, 9, 5618, 2003. 

Luo, D., A new solution for improving gene delivery, Trends Biotechnol., 22, 101, 2004. 

Luo, D. and Saltzman, W. M., Synthetic DNA delivery systems, Nat. Biotechnol., 18, 33, 2000. 

Luo, D. et al., Poly(ethylene glycol)-conjugated PAMAM dendrimer for biocompatible, high-efficiency DNA 

delivery, Macromolecules, 35, 3456, 2002. 

Maeda, H., The enhanced permeability and retention (EPR) effect in tumor vasculature: The key role of tumorselective 

macromolecular drug targeting, Adv. Enzyme Regul., 41, 189, 2001. 

Malik, N. et al., Dendrimers: Relationship between structure and biocompatibility in vitro, and preliminary 

studies on the biodistribution of 125I-labelled polyamidoamine dendrimers in vivo, J. Controlled 

Release, 65, 133, 2000. 

Mannisto, M. et al., Structure-activity relationships of poly(L-lysines): Effects of pegylation and molecular 

shape on physicochemical and biological properties in gene delivery, J. Controlled Release, 83(1), 

169, 2002. 

Maruyama, K., In vivo targeting by liposomes, Biol. Pharm. Bull., 23, 791, 2000. 


548 


Massing, U. and Fuxius, S., Liposomal formulations of anticancer drugs: Selectivity and effectiveness, Drug 

Resist. Updat., 3, 171, 2000. 

McNeil, S. E., Nanotechnology for the biologist, J. Leukoc. Biol., 78, 1–10, 2005. 

Miller, T. M., Kwock, E. W. K., and Neenan, T. X., Macromolecules, 25, 3143–3148, 1992. 

Moghimi, S. M. and Szebeni, J., Stealth liposomes and long circulating nanoparticles: Critical issues in 

pharmacokinetics, opsonization and protein-binding properties, Prog. Lipid Res., 42, 463, 2003. 

NCI Cancer Bulletin, National Cancer Institute, U.S. Department of Health and Human Services, National 

Institutes of Health, NIH Publication, Press Office No. (301) pp. 496–6641, Sep., (http://www.cancer. 

gov), 2004. 

NCI Cancer Bulletin, National Cancer Institute, U.S. Department of Health and Human Services, National 

Institutes of Health, NIH Publication No. 05-5498, Feb., vol. 2, (http://www.cancer.gov), 2005. 

Neerman, M. F. et al., Reduction of drug toxicity using dendrimers based on melamine, Mol. Pharmacol., 1, 

390, 2004. 

Newkome, G. R., Moorefield, C. N., and Vogtle, F., Dendritic Molecules: Concepts, Syntheses, Perspectives, 

VCH, Weinheim, 1996. 

Norsten, T. and Branda, N., A starburst porphyrin polymer: A first generation dendrimer, Chem. Commun., 

1257, 1998. 

Ooya, T., Lee, J., and Park, K., Effects of ethylene glycol-based graft, star-shaped, and dendritic polymers on 

solubilization and controlled release of paclitaxel, J. Controlled Release, 93, 121, 2003. 

Padilla, D. J. et al., Polyester dendritic systems for drug delivery applications: In vitro and in vivo evaluation, 

Bioconjugate Chem., 13, 453, 2002. 

Papahadjopoulus, D. et al., Sterically stabilized liposomes improvement in pharmacokinetics and anti tumour 

therapeutic efficacy, Proc. Natl Acad. Sci., 88, 11460, 1991. 

Peppas, N. A. et al., Polyethylene glycol containing hydrogels in drug delivery, J. Controlled Release, 62, 81, 

1999. 

Phillips, W. T. et al., Polyethylene glycol modified liposomes encapsulated hemoglobin: A long circulating red 

blood cell substitute, J. Pharmacol. Exp. Ther., 288, 665, 1999. 

Qualmann, B. et al., Synthesis of boron-rich lysine dendrimers as protein labels in electron microscopy, 


Qualmann, B. et al., Electron spectroscopic imaging of antigens by reaction with boronated antibodies, 

J. Microsc., 183, 69, 1996. 

Quintana, A. et al., Design and function of a dendrimer-based therapeutic nanodevice targeted to tumor cells 

through the folate receptor, Pharm. Res., 19, 1310, 2002. 

Reuter, J. D. et al., Inhibition of viral adhesion and infection by sialic-acid-conjugated dendritic polymers, 

Bioconjugate Chem., 10, 271, 1999. 

Reynolds, J. E. F., Martindale Extra Pharmacopoeia, 31st ed., Royal Pharmacol Society, London, 1996. 

Sahoo, S. K. and Labhasetwar, V., Nanotech approaches to drug delivery and imaging, Drug Discov. Today,8, 

2003, www.drugdiscoverytoday.com. 

Shukla, S. et al., Synthesis and biological evaluation of folate receptor-targeted boronated PAMAM dendrimers 

as potential agents for neutron capture therapy, Bioconjugate Chem., 14, 158, 2003. 

Sledge, G. and Miller, K., Exploiting the hallmarks of cancer: The future conquest of breast cancer, Eur. 

J. Cancer, 39, 1668, 2003. 

Smith, D. K. and Diederich, F. N., Functional dendrimers: Unique biological mimics, Chem. Eur. J., 4, 1353, 

1998. 

Sparks H., How miniature radiation detectors will keep astronauts safe in deep space. http://www.space.com/ 

businesstechnology/technology/radiation_nanobots_020717.html, 2002. 

Storm, G., Belliot, S. O., Daemen, T., and Lasic, D., Surface modification of nanoparticles to oppose uptake by 

the mononuclear phagocyte system, Adv. Drug Deliv. Rev., 17, 31, 1995. 

Suzawa, T. et al., Enhanced tumor cell selectivity of adriamycin-monoclonal antibody conjugate via a 

poly(ethylene-glycol)-based cleavable linker, J. Controlled Release, 9, 229, 2002. 

Tomalia, D. A., Naylor, A. M., and Goddard, W. A., Starburst dendrimers: Molecular-level control of size, 

shape, surface chemistry, topology, and flexibility from atoms to macroscopic matter, Angew. Chem. 

Int. Ed. Engl., 29, 138, 1990. 

Toyokuni, T., Hakomori, S. I., and Singhal, A. K., Bioorg. Med. Chem., 2, 1119, 1994. 



Veronese, F. M. et al., Preparation, physicochemical and pharmcokinetic characterisation of methoxy polyethylene 

glycol derivatized superoxide dismutase, J. Controlled Release, 10, 145, 1989. 

Vincent, L. et al., Efficacy of dendrimer-mediated angiostatin and TIMP-2 gene delivery on inhibition of tumor 

growth and angiogenesis: In vitro and in vivo studies, Int. J. Cancer, 10, 419, 2003. 

Voit, B., Turner, S. R., and Mourey, T., Hyperbranched polyesters: 1. All aromatic hyperbranched polyesters 

with phenol and acetate end groups; synthesis and characterization, Macromolecules, 26, 4617, 1993. 

Wooley, K. L., Hawker, C. J., and Frechet, J. M. J., A branched-monomer approach for the rapid synthesis of 

dendrimers, Angew. Chem. Int. Ed. Engl., 33, 82, 1994. 

Zalipsky, S., Chemistry of polyethylene glycol conjugates with biologically active molecules, Adv. Drug 

Deliv. Rev., 16, 157, 1995. 

Zeng, F. and Zimmerman, S. C., Dendrimers in supramolecular chemistry: From molecular recognition to selfassembly, 

Chem. Rev., 97, 1681, 1997. 

Zhang, G. D. et al., Polyion complex micelles entrapping cationic dendrimer porphyrin: Effective photosensitizer 

for photodynamic therapy of cancer, J. Controlled Release, 93, 141, 2003. 


28 

CONTENTS 

Dendrimer Nanocomposites 


Lajos P. Balogh and Mohamed K. Khan 

28.1 Composite Nanodevices .................................................................................................... 552 

28.1.1 Introduction ......................................................................................................... 552 

28.1.2 Dendrimers .......................................................................................................... 555 

28.1.3 Toxicity................................................................................................................ 558 

28.1.4 Gold and Radioactive Gold................................................................................. 558 

28.1.5 Dendrimer Nanocomposites ................................................................................ 559 

28.1.5.1 Synthesis of Gold/Dendrimer Nanocomposites................................. 561 

28.1.5.2 Characterization of Dendrimer Hosts and Gold 

Nanocomposites Used in Biodistribution Experiments..................... 562 

28.1.5.3 Fabrication of Nanocomposite Devices for Biologic 

Experiments ........................................................................................ 562 

28.1.5.4 Synthesis of Tritium Labeled Dendrimers......................................... 564 

28.1.5.5 Folate-Targeted Devices .................................................................... 564 

28.1.5.6 Gold Composite Nanodevice Synthesis............................................. 564 

28.1.5.7 Gold in the Nanocomposite Particles After Radiation 

Polymerization.................................................................................... 569 

28.1.5.8 Visualization of Gold/PAMAM Nanocomposite Uptake in 

Tumor Cells ........................................................................................ 569 

28.2 Biology .............................................................................................................................. 571 

28.2.1 Targeting and Nanodevice Delivery ................................................................... 572 

28.2.1.1 Quantitative In Vivo Biodistribution of Dendrimer and Gold 

Composite Nanodevices (CND)......................................................... 573 

28.2.1.2 Related Techniques in the Literature................................................. 573 

28.2.1.3 The Instrumental Neutron Activation Analysis (INAA) Method ..... 574 

28.2.1.4 Determination of Gold Content Using INAA—Precision 

and Sensitivity .................................................................................... 575 

28.2.1.5 Biodistribution of Positive and Neutral Surface Generation Five 

Dendrimers—Non-Specific Uptake ................................................... 575 

28.2.1.6 Biodistribution of 5 nm Gold Nanocomposites Compared with 

Dendrimer Templates ......................................................................... 577 

28.2.1.7 Biologic Differences Between Dendrimers and Gold Composite 

Nanodevices ....................................................................................... 577 

28.2.1.8 The Importance of Size on Nanocomposite Biodistribution............. 579 

28.2.1.9 Biodistribution Summary ................................................................... 580 

28.2.2 Toxicity Studies................................................................................................... 580 


551

552 

28.2.2.1 Long-Term Toxicity Studies of Tritiated 

Dendrimer Nanoparticles ................................................................... 580 

28.2.2.2 Long-Term Toxicity of {Au(0)} Gold Composite 

Nanodevices ....................................................................................... 582 

28.2.2.3 Treatment of Tumors in Mouse Model Systems ............................... 582 


Definitions .................................................................................................................................... 585 

References .................................................................................................................................... 586 

28.1 COMPOSITE NANODEVICES 



Every year, the United States alone reports more than half a million cancer-related deaths and 

approximately 1.3 million new cases. 3 Cancer is customarily treated with drugs (extracts, synthetic 

molecules, and/or biologic materials, e.g., proteins, antibodies, etc.). In an ideal case, the intervening 

medication should only act at the disease site and influence only those mechanisms that are 

causing the illness. Most of the unwanted side-effects are due to therapeutic materials that arrive at 

unwelcome sites (cells, tissues, organs) and influence normal mechanisms in a wrong way. 

Targeted drug delivery (without side effects) has been the ultimate goal of the treatment of diseases. 

Living objects have developed mechanisms to interact and deal with molecules and organic or 

inorganic materials (particles) introduced in the blood stream; there are filter organs and defense 

systems to eliminate the unwanted ones and keep those that are useful. 

Most of the anti-cancer drugs are cytotoxic—they kill cells. To use these in a clinical setting, it 

is mandatory to minimize the side effects, i.e., eliminate cancer cells without compromising the 

normal operation of other cells. The ongoing revolution in nanoscience and nanotechnology offers 

novel ways to treat diseases. These man-made complex objects provide an opportunity to better 

understand how the body works and simultaneously offer new ways of imaging and treatment. 

Targeted nanodevices open up many new avenues for therapy and diagnosis of cancer. 

Understanding biodistribution is a key to success for any medication. Biodistribution is 

determined by the interactions between the properties of the medicine and the living biologic 

system. Therefore, to ensure that a therapeutic material has only one concerted action at the 

molecular level, molecules/particles/devices with identical properties are needed. Mixtures of 

materials will have fractions with different properties, each with an individual biodistribution (of 

which only an envelope is observed). Different fractions with differing properties of biodistribution 

can lead to different actions. In simple terms, only identical particles will have a welldefined 

biodistribution. 

Because of these reasons, such materials need to be prepared that are composed of identical 

molecules/objects, presenting identical properties. In other words, only identical nanodevices with 

well-defined properties will have a defined biodistribution, making homogeneity (or, at least, 

narrow distribution) a fundamental requirement for a successful and reproducible targeted delivery. 

Nanoscience and nanotechnology encompass a broad range of areas with different degrees of 

maturity, namely synthesis and controlled formation of nanostructures with a variety of specific 

functionalities that are enabled by a widespread availability of tools to observe, characterize, and 

manipulate at the nanoscale level. This is a cyclic and synergistic process; improved characterization 

(new methods, higher spatial resolution, and higher sensitivity) leads to a better 

understanding of composition–structure–property relations that motivates better nanoparticle and 

nanostructure fabrication and allows better control of size and placement that, in turn, stimulates 

improved characterization. 


Dendrimer Nanocomposites for Cancer Therapy 553 

Properties of nanosized particles depend not only on composition, but also on size (diameter, 

surface, volume), shape, structure, architecture, flexibility, conformations, etc. Materials composed 

of nanoparticles are characterized not only by their individual properties, but also by the distribution 

of those individual properties over a very large number of particles. To employ nanodevices 

(nanosized multifunctional particles) in therapy and imaging, it is necessary to clearly understand 

the relevant properties of the materials used, how the given biologic system works, and how 

nanodevices interact with the biologic system. 

To ensure the existence of these properties, appropriate characterization methods must be used 

that provide information about relevant properties and their distributions, including nanoparticle 

composition, size, shape, surface, and interior properties. At the nanoscale (1–100 nm), the surfaceto-volume 

ratios are very large, and surface forces (Coulomb interactions, hydrogen bonding, van 

der Waals forces) greatly determine the characteristics of these devices and control the interactions 

of these devices with biological systems. 

Reproducibility of nanoparticle properties is necessary to observe a reproducible biodistribution 

that is key to successful therapy. (Presently, most knowledge of nanoparticles is based on 

demonstrations of concept.) 

Properties and biologic behavior of nanodevices are not yet completely understood, and 

detailed chemical–material science–physics–biology studies are being performed to generate 

more materials with appropriate properties and a better understanding in the future. Empiric 

studies are necessary to uncover any unexpected properties of nanodevices. 

Radiotherapy has been used for decades as standard cancer therapy in almost all forms of 

cancers with varying degrees of success. One of the difficult challenges with radiation therapy is 

the delivery of a lethal dose of radiation to the tumor while leaving the surrounding normal tissue 

unharmed. Radioisotopes are often practically insoluble in water or body fluids, and they are rarely 

used as elements. Typically radioisotopes are attached to an appropriate carrier molecule. Tumordirected 

antibody (or peptide) radiation therapy has been used to achieve this goal with limited 

success. In this technique, the radiation emitter is attached to the targeting antibody or peptide, 

often by a cage-type carrier, that is linked to the targeting molecule. The caged carrier can carry 

only a limited number of radioactive atoms/molecules and usually requires a linker to prevent 

destruction of the targeting molecule. This creates a fundamental limitation for the dose that can be 

delivered per targeting molecule that may limit the ability to successfully treat solid tumors with 

radioactive antibody approaches. For any targeted dose delivery approach, it has to be demonstrated 

that the treatment will be effective if the dose is directly delivered to the tumor 

Composites are physical mixtures or two or more components that display improved properties 

in one or more areas as compared to their individual bulk constituents. They are widely used in 

many diverse industries and form the basics of engineering plastics, structural adhesives, and 

matrices. Nanocomposites can be defined as having at least one component with at least one of 

their dimensions less than 100 nm. Therefore, nanocomposite properties will be strongly influenced 

by extensive interfacial interactions between the components at the interphase, and the rule of 

mixtures may fail to provide estimates of nanocomposite properties. 

Guest–host nanodevices composed of therapeutic (or imaging) materials carried in micelles, 

liposomes, polymeric nanocarriers, dendrimers, nanoemulsions, nanocomposites, etc., are a wellstudied 

class of multifunctional nanomaterials with several potential medical uses, including 

targeted drug-delivery, cancer imaging, and therapy. Typically, the common feature in these 

structures is that the carried material is not bound to the carrier with covalent bonds, but it is 

constrained by physical forces. Composition, 3-D structure, and surface chemistry of the carrier are 

critical parameters for targeted delivery. Charge and size of the nanodevices determine their in vivo 

biodistribution, and thereby, the efficacy for imaging and therapies. 

Composite nanodevices possess chemical and physical (spectral) properties, both the inorganic 

and the organic components, and physical and biologic interactions of the nanoparticles with each 

other and with the environment (solubility, biocompatibility, etc.) are dominated by the size, shape, 


554 


and contact surface of the host. As a consequence, in hybrid composite nanodevices, a specific 

nanoparticle property (e.g., radioactivity belonging to the inorganic guest) can be manipulated as if 

it were a property of the organic host (e.g., dendrimer). 

Targeted composite nanodevices are constructed by covalently binding targeting moieties to 

the organic template that dominates the surface of the nanocomposite. The surface can be chemically 

modified with various substituents, including proteins to achieve the necessary organ/tissue 

specificity and create targeted composite nanodevices. The interaction of these nanodevices with 

the biologic environment (solubility, biocompatibility) is controlled by the contact surface of the 

host molecule, i.e., by their size and surface properties. The separation of overall properties from 

the surface chemistry (targeting, charge, and size) is a unique advantage in the optimization of these 

nanodevices because properties of the components can be individually selected and optimized. 4 

Nearly monodisperse gold/dendrimer nanocomposites can be synthesized in various predetermined 

sizes, and their interaction with biological objects may be adjusted by modifying their surface 

properties. They can readily penetrate cells and are easy to observe by various methods. 5 Composite 

nanoparticles with either cationic, anionic, neutral, lipophilic, lipophobic, or mixed surfaces can be 

created by encapsulating passive or active therapeutic agent(s) that adsorb or emit radiations, 

respectively. Guests in radioactive nanocomposite devices may be metals or other compounds 

containing isotopes of medical importance permitting imaging labels, (e.g., 125 I, 3 H) or radiation 

therapy, for example, 198 Au that can deliver b-radiation to tumors. 

Using targeted nanodevices allows one to deliver various functionalities via a single biohybrid 

nanoparticle, making these nanodevices a potentially attractive tool in radiation medicine. An 

analysis of tumor type should allow for the selection and rapid immobilization of the appropriate 

radioactive isotope in the pre-made template that can then be efficiently delivered to specifically kill 

tumor cells with minimal collateral damage. The charge, size (i.e., surface chemistry), composition, 

and 3-D structure of these nanoparticles are critical in determining their in vivo biodistribution, and 

therefore, the efficacy of nanodevice imaging and therapies. 

Some of the fundamental problems that can be overcome by the use of targeted composite 

nanodevices are solubility and dose delivered. First, many radioisotopes are often practically 

insoluble in water or body fluids, but they may become readily available in nanocomposites, i.e., 

in different sizes, contents, and surfaces. Second, use of composite nanodevices permits at least a 

log-fold higher delivery of radioactivity than that available with current radioactive antibody 

therapies. As opposed to labeled antibodies, at least 1–1000 radioisotope nuclei can be encapsulated 

in one nanocomposite device 6. Therefore, radioactive nanocomposite particles will be able to 

deliver therapeutic dose amounts to tumor cells. (Recent data indicate that even one radioactive 

alpha-particle decay could trigger cell death if it occurs in the cell.) 7 

The specific advantage of the targeted dose delivery by composite nanodevices approach over 

intracellular drug delivery or simple diffusion of drug into cells is that cell surface delivery is 

sufficient and no internalization is needed. As the specific activity of one composite { 198 Au} 

nanodevice is extremely small (in the range of 1E-15 Ci/nanoparticles, i.e., 3.7E-25 Bq/nanoparticle), 

single particles delivering b-radiation at the cell surface (low-level non-specific uptake 

by T-cells) are harmless, but concentrations above a certain critical level (i.e., when large number 

of particles are delivered to specific cells or tissues) can be therapeutic. 

Practically, for any targeted delivery approach, it has to be demonstrated that the treatment is 

effective if the therapeutic material accumulates in the tumor. These conceptual experiments must 

be followed by biodistribution studies and the optimization of devices. 

In a proof-of-concept research, { 198 Au} nanodevices were tested, but conceptually, a number of 

other nuclides with the desired specific activity could be used (allowing the use of gamma, beta, 

alpha, or combinations of these radiations in the future). As gold is non-toxic and the hot and cold 

gold isotopes are chemically identical, biodistribution studies may conveniently be done with 

nanocomposites containing non-radioactive gold. 



28.1.2 DENDRIMERS 

The term dendrimer may refer to a mathematical structure, a chemical structure of a single 

molecule of a certain class of macromolecules, or to a certain material (usually a product of a 

specific synthesis procedure) 8,9 (Figure 28.1). A dendrimer material is typically a mixture of very 

similar polymer molecules. 

A number of characteristic general properties of dendrimers can be concluded from the 

dendritic structure itself as a result of symmetry and maximal degree of branching. These characteristics 

are the function of the core structure, the connector length, the degree of branching, etc., 

and they appear in all the families (for example, the tendency of dendrons to move in a concerted 

fashion). 10–13 

Dendrimer molecules are symmetric molecules with possible molecular masses of up to several 

million daltons (i.e., dendritic structures built from atoms) that contain connectors and branching 

units composed of repetitive shell structures around a core following a predefined molecular motif. 

Dendrimer molecules always have specific chemistry and specific properties. 

There are many ways to construct polymer molecules with dendritic architecture, depending on 

which part of the macromolecule is dendritic (and what different authors consider to be a 

dendrimer). To date, more than 200 various dendritic structures have been reported in the literature. 

The major synthesis strategies are the divergent route that leads to higher yields with lower 

dendritic purity. 14 Convergent strategies 15 usually result in more uniform molecules but with 

lower overall yields (a great number of other strategies have been also reported). Dendrimer 

molecules are well-defined and highly symmetrical molecules containing a large number of regularly 

spaced internal and external functional groups. (Dendritic molecules are built to maximize 

symmetric branching.) As a result, the interior of a dendrimer may considerably differ from its 

exterior and may be hydrophilic or hydrophobic, depending on the design and synthesis route. 

Constructing a dendritic network from atoms necessarily invokes specific properties because 

the atoms themselves have specific properties. Dendrimer families contain a certain group of 

molecular motifs (PAMAM, PPI, polylysine dendrimers, polyether dendrimers, polyacetylene 

dendrimers, etc.). They are synthesized on different synthetic principles and use different routes; 

therefore, they have different chemistries. Dendrimer families may be very different. 

Generations and their subclasses. Within any families there are low, high, and middle generation 

dendrimers. Molecules of low generation dendrimers (LGD) structurally are relatively small 

organic molecules with fully flexible loose networks and with a few interactive sites (e.g., the 

FIGURE 28.1 (a) Dendritic structure (three-functional core, bifunctional branching, persistent angles) and a 

computer model of an ideal PAMAM dendrimer molecule (b). (The left panel has been reproduced from 

Bielinska, A., Eichman, J. D., Lee, I., Baker, J. R., and Balogh, L., J. Nanopart. Res., 4, 395–403, 2002. With 

permission.) 


556 


chemical accessibility of the functions are roughly equal). High generation dendrimers (HGD) are 

essentially highly charged organic nanoparticles where the accessibility of surface groups is much 

higher than of the interior. These organic nanoparticles have a well-defined generational size and a 

persistent surface, properties that may be quite different from its interior. Medium generation 

dendrimers (MGD) are transitional and may display properties of both groups. 

A dendrimer material is typically a mixture of very similar polymer molecules as only perfect 

reactions (complete conversion with 100% yield) or perfect purification would lead to perfect 

structures in the consecutive synthesis steps. Consequently, practical dendrimer materials are 

oligomers or macromolecules with narrow polydispersity, and they always contain molecules 

with imperfect structures. However, the high level of synthetic control and effective purification 

between the synthesis steps allows the synthesis of defined polymers with very low polydispersities. 

Material properties are usually different for different generations; for example, in viscosity 

measurements, LGD materials can be described as draining networks, whereas HGD materials 

behave as ideal Newtonian fluids. There may be measurable differences between batches of individually 

manufactured or synthesized dendrimer materials; therefore, one must be very careful to 

generalize experimental results to the class of dendrimers. Dendritic properties of a given material 

can neither be assessed nor evaluated without taking into account the family, generation, composition, 

various properties, and their distribution (Figure 28.2). 

Polyionic water-soluble dendrimers are those where tertiary nitrogen provides the branching. 

Poly(amidoamine) (PAMAM) dendrimers contain beta-alanine subunits and can be synthesized to 

be biofriendly. 16–19 They are well characterized and are also commercially available. PAMAM 

dendrimer molecules undergo changes in size, shape, and flexibility as a function of increasing 

generations (from a small branching molecule at generation 0–2 (LGD) through flexible macromolecules 

(MGD: G3, G4, G5–G6) to dense organic particles behaving as hard spheres at 

generations seven and above) (Figure 28.2). Branching structure of MGD and HGD PAMAMs 

can efficiently entrap therapeutic molecules. 20–22 They have been used as delivery vehicles for 

oligonucleotides and antisense oligonucleotides and as probes for oligonucleotide arrays. 23,24 

PAMAM dendrimers are not broken down by enzymes in vivo and generally are removed from 

the bloodstream by the filter organs. 25–27 

Diversities that exist in dendrimer materials are of generational, skeletal, and/or substitutional 

origin. It is common that traces of various generations are present in materials (Figure 28.3a, peaks 

A–B–C in SEC-MALLS); the individual molecules may or may not be perfectly dendritic (see: 

skeletal diversity), and less than complete substitution of the end groups results in slightly different 

substitutions on the individual molecules (because of the random nature of chemical reactions, i.e., 

FIGURE 28.2 Size and shape of ethylenediamine core PAMAM dendrimers as a function of generation (fully 

protonated structures in water, i.e., maximally extended conformations). Dendritic polymers provide a multifunctional 

surface with a variable size and tunable compatibility. (Reproduced from Bielinska, A., Eichman, J. 

D., Lee, I., Baker, J. R., and Balogh, L., J. Nanopart. Res., 4, 395–403, 2002. With permission.) 



Relative scale 

1.2 

1 

0.8 

0.6 

0.4 

0.2 

0 

B 

C A 

LS 

LV 

RI 

–0.2 

30 31 32 33 34 35 36 37 38 

(a) 

Time (min) 

substitutional diversity). 28 The most important properties of dendrimers are mass and charge 

(Figure 28.3). In case of HGD (e.g., GO6 in the case of PAMAMs) that are dense-packed, 

contribution to the surface may be solely due to terminal groups. In the more open structures of 

lower generations, other functional groups also contribute in an increasing fashion as the generational 

number decreases. 

Dendrimers have a very rich synthetic chemistry describing how molecules with dendritic 

architectures can be synthesized and/or modified (to date, more than 5000 papers have been 

published on dendrimers). 8,9 Large dendritic structures may be synthesized by reacting dendrimers 

as building blocks (e.g., as core and shell reagents) with each other (Figure 28.4). These molecules 

have been termed tecto-dendrimers. 21,29,30 

Among other means, PAMAM dendrimers have been proposed as versatile vehicles of drug 

transport to specific organs and tissues because of their unique external and internal functionalities 

that can be altered according to need. These spherical macromolecules have modifiable surface 

functionalities offering a versatile mode to covalently attach drugs, diagnostic/imaging modules, 

and/or targeting moieties, and they allow the control of effective surface charge. However, the 

expensive starting materials, the required multistep synthesis, and subsequent characterization of 

the multifunctional nanodevices often present a serious challenge. 31 This complexity may be 

Relative scale 

60 

50 

40 

30 

20 

10 

0 

10.54 min 

–10 

0 5 10 15 20 

(b) 

Time/min 

E5.NH2 

25 30 35 

FIGURE 28.3 (a) Size exclusion chromatogram (SEC-MALLS) showing mass (MnZ28,940) and polydispersity 

(PDIZ1.067) A: G4 trailing generation; B: G5 main component; C: dimers of G5 present in the 

material; and (b) a capillary electropherogram (CE) illustrating the almost identical charge/mass ratio for an 

actual ethylenediamine core primary amine-terminated generation five PAMAM dendrimer material. 

G y 

Core reagent 

(NH 2 ) 

− 

G x 

Shell reagent 

(excess) 

(1) Self assembly 

(Equilibration) 

CO2H (2) Covalent bond 

formation 

G x 

G x 

G x 

G y 

G x 

G x 

G x 

Generation 1 core-shell 

Tecto(dendrimer) 

Gz 

(NH 2 ) 

G z 

G z 

G z 

G z 

G x 

G x 

G z 

Gz 

G x 

G y 

G x 

G z 

G z 

G x 

G x 

G z 

G z 

G z 

G z 

Generation 1 core-shell 

Tecto(dendrimer) 

FIGURE 28.4 Scheme of tecto-(dendrimer) synthesis. The method is used to make large templates and supramolecular 

nanoparticles for multi-functional composite nanodevice fabrication. 


558 

considerably simplified using hybrid composite nanoparticles templated by monodisperse dendrimers 


Polyionic dendrimers are close to being optimal templates for these composite nanoparticles 

because the better the templates, the more homogenous the resulting composite particles become. 

PAMAM and PPI can interact with a wide variety of ions and in situ synthesized compounds. 32–38 

These tertiary nitrogens act as focal points for pre-orientation either of cations to coordinatively 

bind through donor–acceptor interactions of the lone electron pair or of anions that can electrostatically 

bind to the positive charge of the protonated –NZligands. Multiple binding within the 

large molecule (a nanoscopic size pseudo-phase) also contributes to the unusually strong binding 

and high binding capacity. 39,40 Therefore, using PAMAM dendrimers as templates, the pre-orientation 

and dispersion of the inorganic components are controlled by the appropriate dendritic 

polymer host 32,34,35 and the size, size-distribution, shape, and surface of the resulting composite 

particle can be designed. Countless combinations are possible between metal cations or anions 

and dendrimers, but examples will be limited to aqueous solutions of generation (g4–g6) ethylenediamine 

core poly(amidoamine) (PAMAM) dendrimers and gold in the form of complex 

[AuCl 4] K (tetrachloroaurate) anions. 41,42 

28.1.3 TOXICITY 

+MY 

+X 

D [(MY) n –D] {(MX) n –D} 

–Y 

Dendrimer 

Dendrimer biocompatibility and toxicity have been recently reviewed, and details can be found in 

the literature. 16 Briefly,ininvitroexperiments, it was observed 43 that amine-terminated G5 

PAMAM dendrimers decreased the integrity of the cell membrane and allowed the diffusion of 

cytosolic proteins out of the cell and dye molecules into the cell. Although neither G5 amine- nor 

acetamide-terminated PAMAM dendrimers were cytotoxic up to 500 nM concentration, the dose 

dependent release of the cytoplasmic proteins lactate dehydrogenase (LDH) and luciferase (Luc) 

indicated the presence of holes. The induction of permeability caused by the amine-terminated 

dendrimer was not permanent, and leaking of cytosolic enzymes returned to normal levels upon 

removal of the dendrimers. Diffusion of dendrimers through holes is sufficient to explain the nonspecific 

uptake of positively charged PAMAM dendrimers into cells and is consistent with the lack 

of uptake of neutral surface PAMAM dendrimers. 

28.1.4 GOLD AND RADIOACTIVE GOLD 


Complex/Salt Complex/Sal t 

Nanocomposite 

FIGURE 28.5 General scheme of dendrimer nanocomposite synthesis through complex formation followed 

by reactive encapsulation. (Reproduced from Bielinska, A., Eichman, J. D., Lee, I., Baker, J. R., and Balogh, 

L., J. Nanopart. Res., 4, 395–403, 2002. With permission.) 

Noble metals, especially gold, have many applications in optical and biological sciences as a result 

of their stability. Metallic gold is usually produced by reduction using Na-citrate, citric acid, 



hydrazine, hydrogen, or electrochemical deposition, and the color of these particles is the function 

of size and shape of the metal domains formed. 44 

Gold-based therapeutic agents have been recently reviewed. 45 Gold nanoparticles 46 and nanoshells 

47–49 have recently been successfully tested for photothermal therapy of cancer. Radioactive 

gold has been approved by FDA for radiation treatment and the use of 198 Au colloids for cancer 

therapy has already been attempted in the 1950s. 45,50–52 Radioactive 198 Au is used in interstitial 

brachytherapy for treatment of prostate cancer, initially in the form of gold colloids (micron-sized 

particles), 53 and later with seed technique. 54 

With biopsy results approaching a negative rate of 80%, and at five years, a cancer specific 

survival of 100% for Stages A and B1, 90% for Stage B2, and 76% for Stage C, this form of 

treatment offers an effective and well-tolerated alternative mode of therapy for patients with 

localized prostate cancer. Present evaluation of the accumulated Au(0) related biodistribution 

data is almost impossible because of the partial or full lack of reliable particle characterization 

describing size and/or surface properties 51,52 and literature thereof. 

Use of radioactive gold cluster immunoconjugates (containing 11–33 Au atoms) was reported 

in the early 1990s, but because of the cumbersome preparation and insufficient delivered dose, the 

research was discontinued. 55 Recently, 30–34 nm multifunctional, TNF-conjugated gold nanoparticles 

were prepared and tested for targeting solid tumors in mice. 36,41,56,57 

28.1.5 DENDRIMER NANOCOMPOSITES 

Dendrimer templated nanocomposites are synthesized by reactive encapsulation. 58,59 According to 

the general concept of DNCs, 32 precursors to a desired product are pre-organized by an appropriately 

selected dendrimer in the first step. This pre-organization is followed by in situ chemical 

reaction(s) or physical treatment (irradiation, etc.) that generates reaction products immobilized in 

the nanoscopic size dendrimer molecules (Figure 28.5). This procedure yields dispersed small 

domains of guest molecules that are integrated with the dendrimer molecule(s) without creating 

covalent bonds between the dendrimer and the topologically entrapped matter. Dendrimer 

templated nanocomposites may be constructed with flexible architectures and in 

variable compositions. 

Structure of nanocomposite particles was found to be the function of the dendrimer structure 

and surface groups as well as the formation mechanism and the involved chemistry. Based on the 

locality of the guest atoms, three different types of single nanocomposite architectures have been 

identified such as internal (I), external (E), and mixed (M) type nanocomposites where the inorganic 

matter either entrapped in the interior, in the exterior, or in both part of the dendrimer, respectively. 

Dominant properties of the composite nanoparticles approach the dendrimer properties with 

decreasing M/D ratios (i.e., with the molar ratio of the inorganic reactants and the organic 

template). Stable internal structures form when medium and high generation polyionic dendrimers 

with a weak binding periphery are employed with proper chemistry and at relatively low metal/dendrimer 

ratios. (In this case, the inorganic domains are located in the interior of the dendrimer 

templates.) Single composite nanoparticles may be used as building blocks in the synthesis of 

nanostructured materials and devices. Details of experimental procedures and analytical techniques 

can be found in previous studies. 41 

Dendrimer nanocomposites can be constructed in precise sizes and well-defined surfaces with 

various covalently attached targeting moieties and with variable content. DNCs display stability 

and robustness in addition to properties represented by their individual components as a result of the 

molecular level’s mixing of uniform macromolecules and nanoparticles. The unique advantage of 

this approach is that host and guest components can be separately selected, developed, and 

optimized. As a result of various synthetic options, the interior and/or the exterior of the host 

can be cationic, anionic, or non-ionic, depending on their termini and interior functionalities and 

the pH. 


560 


FIGURE 28.6 Scheme of assembling templates and synthesis of nanocomposites of well-defined sizes and 

surfaces (a) and TEM image of an {Au} nanocomposite single particle and a dimer of {Au-PAMAM} n multiparticle 

(b). (Reproduced from Bielinska, A., Eichman, J. D., Lee, I., Baker, J. R., and Balogh, L., J. Nanopart. 

Res., 4, 395–403, 2002. With permission.) 

Preparation of modular composite structures is possible either (a) by the synthesis of modular 

templates from simple dendrimer molecules (i.e., using tecto-dendrimers), or alternatively, (b) 

synthesizing of modular nanocomposites from single nanocomposite units. Variation of nanocomposite 

size is possible by using larger templates such as tecto-dendrimers or by combining existing 

nanocomposite particles. Alternatively, the size of nanoparticles may also be varied by reacting the 

surface of premade nanoparticles (Figure 28.6a and Figure 28.6b). Stable internal composite 

structures cannot be formed using linear macromolecules, only from dendrimers. This is the substructure 

that has great potential in combining the properties of organic and inorganic materials for 

biomedical applications. For these nanocomposite structures, the particle surface is supplied by the 

macromolecule; therefore, the solubility and surface compatibility of the nanocomposite particles is 

very similar to the template dendrimer macromolecules (Figure 28.7). These DNC structures are 

most frequently amorphous or nearly amorphous. 60 

Radioisotopes are often practically insoluble in water or body fluids; however, the nanocomposites 

are soluble in different sizes, contents, and surfaces. An analysis of tumor type should allow 

for the selection of a targeter, allowing rapid immobilization of the appropriate radioactive isotope 

in the pre-made template. Targeted composite nanoparticles can then be efficiently delivered to 

specifically kill tumor cells with minimal collateral damage. 

As a basis for a future targeted 198 Au-based image-guided therapy, the effects of size and 

surface charge on the in vivo biodistribution, excretion, and toxicity of gold/dendrimer 

FIGURE 28.7 (a) Computer simulation of a dendrimer nanocomposite; (b) HRTEM of a gold nanocomposite 

particle containing 14 Au atoms. (Reproduced from Khan, M. K., Nigavekar, S. S., Minc, L. D., Kariapper, M. 

S., Nair, B. M., Lesniak, W. G., and Balogh, L. P., Technol. Cancer Res. Treat. 4(6), 603–613, 2005. With 

permission.) 



nanocomposites in a mouse melanoma tumor model system were studied. The nanodevices tested 

include 5 and 22 nm primary amine surface, 5 nm carboxylate surface, and 11 nm carboxylate 

surface gold/dendrimer nanocomposites that display positive and negative surface charges at biological 

pH values. Gold nanocomposites offer the high electron density and radioactivity of the 

immobilized gold atoms, and the interactions of the composite particle with the surrounding 

biologic environment are determined by the surface. Biologic fate can be influenced by increasing 

or decreasing the number and/or strength of interacting forces on the nanparticle surface by changing 

the number and character of the dendrimer terminal groups. 

28.1.5.1 Synthesis of Gold/Dendrimer Nanocomposites 

To synthesize gold/dendrimer composite nanodevices in nearly uniform sizes and with close to 

identical properties, polyionic PAMAM dendrimer templates with low (4–10) M/D ratios were 

used. This method leads to well-controlled and uniform composite particles. The product is a 

synthetic dendrimer-dominated material that is water-soluble and stable. These hybrid nanoparticles 

have tunable properties; they can be made to be non-immunogenic, and they are similar in size 

to fundamental proteins present in the blood. 61 

Synthesis of {Au(0)} gold/dendrimer composite nanoparticles from dendrimer templates 

(Figure 28.5) involves two synthesis steps: (a) forming a PAMAM-tetrachloroaurate salt, and (b) 

the transformation of [AuCl4]K through Au 3C ions to Au(0). Hydrolysis of [AuCl4] K is a complex 

process 62 and easily occurs in aqueous solutions of PAMAMs as the nitrogens neutralize HCl, the 

hydrolysis byproduct. Reduction is usually performed either by adding external reducing agents 

(N2H4 or NaBH4) 41,57 or utilizing the reductive nature of terminal groups of PAMAMs. 63 

In the two extreme cases, (depending on the reaction conditions) the dendrimers may either 

encapsulate the colloidal gold, 57 or alternatively, gold nanocrystals may grow in the interior of the 

PAMAM molecules. 6 Reduction of Au 3C ions by the PAMAM itself may be accelerated by 

applying UV-irradiation. 64 The usual size of dendrimer templated nanocomposite particles is 

dZ3–10 nm, depending on the size of the template, the gold/dendrimer ratio, and the type of 

the composite particle. 41,65 Smaller (LGD) PAMAMs provide larger, multiparticle composites 

because of their open structures, whereas for MGD and HGD (G5–G9) PAMAMs and at low 

metal/dendrimer ratios, the composite particles preserve the size of the template. 66,67 

Controlled fabrication of large nanocomposite particles requires large templates. Synthesis of 

positively charged large (dO10 nm) PAMAM dendrimer structures, (e.g., generation 9 PAMAM 

with an MnZ467,000 and dZ11.4 nm) the routinely used divergent route requires several 

months. The multi-step synthesis and purification (altogether 18 step for G9) results in expensive 

(presently $16,500/g), materials. 68 Applying tecto-dendrimers 69–71 as templates is a more favorable 

approach, but tecto-PAMAMs can provide only methylester or carboxylate terminated (i.e., negatively 

charged) composite nanoparticles. 

It has been found that neutron/gamma irradiation also induces elemental gold formation in 

aqueous solutions or gels (submitted for publication). A novel and simple procedure to fabricate 

{( 198 Au(0)n-PAMAM}C positively charged radioactive gold nanocomposites in dZ10–30 nm 

sizes directly from PAMAM tetrachloroaurate salts or dZ5 nm {Au(0) n-PAMAM_E5.NH 2} kC 

nanocomposites by simultaneously utilizing both the neutron and gamma-radiation have been 

devised and developed. This method was used to fabricate positively charged radioactive gold 

nanocomposites in sizes between 10 nm and 30 nm, also to be used as dose delivery agents (Nano 

Letters, 2006, submitted). 

Particle size, mass, and charge of dendrimer nanocomposites may be very similar to their 

templates. Shown in Figure 28.8 is the PAGE electropherogram of different generations of 

amine-terminated dendrimer templates and the corresponding dendrimer nanocomposites. These 

Au-nanocomposites were templated by polycationic (amine terminated) PAMAM dendrimers of 

different generations with the same molar ratio of nitrogen ligand/gold atom and were characterized 


562 

FIGURE 28.8 PAGE electropherograms of several polycationic PAMAM dendrimer templates and their 

corresponding gold-dendrimer nanocomposites. Observe the similarity between the movement patterns of 

the dendrimer templates and the nanocomposites, indicating similar charge/mass ratios for the templates 

and the gold nanocomposites (D nitrogen ligand/Au ratioZ4.5–5.2). (Reproduced from Khan, M. K., Nigavekar, 

S. S., Minc, L. D., Kariapper, M. S., Nair, B. M., Lesniak, W. G., and Balogh, L. P., Technol. Cancer 

Res. Treat. 4(6), 603–613, 2005. With permission.) 

after a prolonged storage, i.e., in thermodynamic equilibrium. The amine-terminated golddendrimer 

nanocomposites and their respective dendrimer templates all display very similar 

migration patterns. (Au-nanocomposites display somewhat lower electrophoretic mobility probably 

because of the slightly higher mass of the composite particles.) 65 

In any case, using a good quality template is essential for success. Templates can ease the 

difficulty of defining size and shape, but the quality of the product will be determined by 

the template’s precision. In short, the product cannot be better than the template and the procedure 

used. In an ideal case, one would have a perfect template and a subsequent perfect procedure to get 

a homogenous and uniformly sized nanomaterial of a predefined shape. Any imperfection in the 

template or the procedure results in a broadening distribution of composition and geometry (and, 

subsequently, of the properties) in the product. 

28.1.5.2 Characterization of Dendrimer Hosts and Gold Nanocomposites Used in 

Biodistribution Experiments 

Determination of fundamental properties (composition, size, and charge) is vital for reproducibility 

prior to initiating any biological experiment. A variety of analytical techniques to were applied to 

investigate the purity, properties, and structural characteristics of dendrimer hosts and host-guest 

nanocomposites, including polyacrylamide gel electrophoresis (PAGE), capillary electrophoresis 

(CE), 72 size exclusion chromatography (SEC), and HPLC, 73,74 UV–visible and fluorescence 

spectrophotometry, acid–base potentiometric titration, mass spectrometry (ESI-MS and MALDI- 

TOF techniques), and various NMR techniques. 65,75,76 (available at www.mrs.org). 77 Gold/dendrimer 

nanocomposites have been characterized by UV–visible and fluorescence spectrophotometry, 

NMR, DLS (dynamic light scattering), and TEM. Particle structure of the neutron-irradiated 

gold nanocomposites have also been studied (after isotope decay) by methods described above 

in order to evaluate the effect of direct activation on the stability of the polymer host. 4,42,61 

28.1.5.3 Fabrication of Nanocomposite Devices for Biologic Experiments 


In these studies, non-targeted and folic acid targeted hosts were synthesized and characterized (i.e., 

dZ5 nm) 3 H labeled PAMAM dendrimers with positive and neutral surface charges in normal and 



TABLE 28.1 

Molecular Characteristics of Polyanionic PAMAM Dendrimer Templates 

Mw a 

Mn b 

Mw b 

Polydispersity b 

PAMAM E3.SAH E5.SAH E5(E3.SAH) 9 

10,109 41,626 147,542 

6300 38,400 152,200 

9790 38960 182,100 

1.022 1.045 1.196 

Theoretical no. of COOH groups a 

32 128 256 

Practical no. of carboxylate groups c 

30.6 113.2 296 

Charge/mass (mol/g) a,d 

K3.18!10 K3 

K3.08!10 K3 

Charge/mass (mol/g) b,c 

K3.32!10 K3 

K2.90!10 K3 

a 

Theoretical molecular weight assuming complete conversions. 

b 

As measured by SEC at pH 2.74. 

c 

Measured by potentiometric titration. 

d As measured by CE. 

K1.74!10 K3 

K1.94!10 K3 

tumor tissues to examine the biodistribution and the toxicity. Then, gold nanocomposites with 

different surface charges and in different sizes were prepared and studied. 

The combined analytical results and properties of the PAMAM dendrimers used as templates 

are presented in Table 28.1 and Table 28.2. Details of template synthesis and characterization are 

described in the literature. 

Zeta potential measurements of {(Au 0 )9-PAMAM_E5.(NH2)110} in aqueous solutions 

confirmed that these nanoparticles are positively charged, indicating that the terminal amines 

can still be protonated after the formation of the nanocomposites. The net charge of amine-terminated 

dendrimers and their Au-nanocomposites was compared by PAGE and proved to be similar to their 

corresponding templates in the pHZ8.3 running buffer. 65 

TABLE 28.2 

Molecular Characteristics of Polycationic PAMAM Dendrimer Templates 

Mn a 

Mn b 

Dendrimer E3.NH2 E5.NH2 

6909 28,826 

6648 27250 c 

Diameter a (nm) 3.6 5.4 

Theoretical no. of NH2 groups a 

32 128 

Total no. of N ligands a 

62 254 

Charge/Mw ratio d (C/g) 8.973 !10 K3 

8.811 !10 K3 

Average no. of –NH2 groups b 

26.3 120 c 

Total average no. of –NZligands b 

54 238 c 

Charge/Mw ratio b (C/g) 8.12 !10 K3 

a 

Theoretical values. 

b 

Practical values determined from titration and GPC measurements. 

c 

Literature data. 

d 

Assuming theoretical values and full protonation. 


8.73 !10 K3

564 

28.1.5.4 Synthesis of Tritium Labeled Dendrimers 

Earlier biodistribution studies of PAMAM dendrimers have been carried out with full generation 

PAMAM dendrimers that were radiolabeled by partially quaternizing the primary amine termini 

using 14 C, containing methyl iodide or 125 I labels. 17,78 Because these radiolabeled macromolecules 

carry permanent positive charges, this labeling procedure results in a mixture of products where the 

labeled molecules are chemically different from the unmodified population of dendrimers. The 

iodine-labeling procedure may lead to data where the distribution of radiolabeled materials is 

different from the overall distribution of the nanoparticles. 

Labeling of dendrimers by 3 H allows proper and accurate quantification of the PAMAM 

derivatives for the true in vivo biodistribution as a function of time because the labeled molecules 

are chemically identical with the unlabeled dendrimers. Regulation of surface charge with simultaneous 

labeling with tritium was achieved via acetylation by partially or fully reacting terminal 

primary amine groups of the PAMAM molecules. This synthetic step reduces the in vivo toxicity 

associated with the polyamine character (the terminal primary amine groups are positively charged 

at biologic pH values) and provides stable labels for the dendrimers. Increasing the degree of 

acetylation decreases the surface net positive charges; the fully acetylated generation five 

PAMAM displays about C5 mV zeta potential as opposed to C40 mV measured for the 

primary amine functional PAMAM_E5.NH2. 

The biodistribution of the fully acetylated dendrimer (NSD, Neutral Surface Dendrimer) and a 

partially acetylated one that has a partially positive surface charge have been compared. (The term 

of positively charged refers to the surface character in dilute aqueous solutions at biologic pH.) 

28.1.5.5 Folate-Targeted Devices 

Briefly, partially acetylated PAMAM was conjugated to folic acid as a targeting agent, and the 

remaining primary amines were capped using tritium labeled acetic anhydride. These conjugates 

were intravenously injected into immunodeficient mice bearing human KB tumors that overexpress 

the folic acid receptor. In contrast to non-targeted polymer, folate-conjugated nanoparticles 

concentrated in the tumor and liver tissue over fours days after administration. The tumor tissue 

localization of the folate-targeted polymer could be attenuated by prior i.v. injection of free folic 

acid. Using fluorescein labels instead of tritium labels, confocal microscopy was also used to 

confirm the internalization of the conjugates into the tumor cells. PAMAM dendrimers substituted 

with folic acid and methotrexate increased its anti-tumor activity and markedly decreased its 

toxicity, allowing therapeutic responses not possible with a free drug. 79 

28.1.5.6 Gold Composite Nanodevice Synthesis 


The {Au} gold/dendrimer composite nanodevices described below have been fabricated using four 

methods: 

1. By reactive encapsulation of gold in commercial (primary amine terminated and 

carboxylate terminated) PAMAM templates; 

2. By performing post-synthetic modificationonthesurfacegroupsofthe{Au}gold 

nanocomposites; 

3. By reactive encapsulation of gold in carboxylate functional PAMAM tecto-dendrimer 

templates; and 

4. By radiation polymerization of {Au} nanocomposites (Table 28.3). 

Method 28A1 fðAu 0 Þ 9:08KPAMAM_E5:ðNH 2Þ 120g C dZ5 nm, for short: fAuð0Þg C dZ5 nm, i.e., dZ 

5 nm amine surface gold/dendrimer nanocomposite (positively charged at pHZ7.4). 



TABLE 28.3 

Summary of Nanodevices and Notations 

Description Full Annotation Acronym 

PSD PAMAM_E5:ðNH2Þ44ðNHCOCH3 Þ66 E5.(NH2)(NHOAc a ) 

NSD PAMAM_E5:ðNHCOCH3 Þ110 E5.OAc a 

Targeted NSD PAMAM_E5:ðNHCOCH3 Þ104ðFAÞ6 E5.OAc a Neutral surface 5 nm gold CND fðAu 

-FA 

0Þ9:08KPAMAM_E5:ðNHCOCH3Þ120g C dZ5 nm fAuð0Þg 0 dZ5 nm 

Positive surface 5 nm gold CND fðAu0Þ9:08KPAMAM_E5:ðNH2Þ120g C dZ5 nm fAuð0Þg C dZ5 nm 

Negative surface 5 nm gold CND fðAu0Þ6:45KPAMAM_E4:5ðCOOHÞ60g K dZ5 nm fAuð0Þg K dZ5 nm 

Negative surface 11 nm gold CND fðAu 0 Þ60:09KðPAMAM_ððE5ÞðE3:ðCOOHÞ296ÞÞ Polymerized PCND b 

Polymerized PCND b 

The dZ5 nm positive nanodevices (Figure 28.7) were fabricated from PAMAM dendrimers 

and HAuCl4 as previously described. 5,36 Briefly, solution of HAuCl4 in methanol was added dropwise 

to a methanol solution of PAMAM dendrimer (MnZ26,000, functionalityZ120), resulting 

in a yellow solution of the tetrachloroaurate salt of the PAMAM dendrimer with a gold/dendrimer 

ratio of Au/DZ9.08), i.e., [PAMAM_E5.(NH2)120)(AuCl4)9.08]. The resulting yellowish solid was 

redissolved in water that then resulted in hydrolysis of the complex and led to nanocomposite 

formation (composite gold contentZ 6.44%). (Positive composite nanodevice, 

PCND)(Figure 28.9). 

Method 28A2 fðAu 0 Þ 6:45KPAMAM_E4:5ðCOOHÞ 60g K dZ5 nm, for short: fAuð0Þg K dZ5 nm, dZ5 nm 

carboxylate surface gold/dendrimer nanocomposite (negatively charged at pHZ7.4) 

Briefly, methanol solution of HAuCl4 was added drop-wise to an aqueous solution of generation 

4.5 PAMAM dendrimer (MnZ26,258, functionalityZ60), resulting in a red solution of the 

correspondent dendrimer nanocomposite (Au/DZ 6.45). (Composite gold contentZ4.41%) 

(Negative composite nanodevice, NgCND). 

Method 28B. Acetylation of the fAuð0Þg C dZ5 nm amine surface material, resulted in 

fðAu 0 Þ 9:08KPAMAM_E5:ðNHOCCH 3Þ 110g 0 dZ5 nm, for short: fAuð0Þg 0 dZ5 nm, an acetamide surface 

dZ5 nm gold/dendrimer nanocomposite that is nearly neutral at pHZ7.4 (neutral composite 

nanodevice, NCND). 

g K dZ11 nm 

fAuð0Þg K dZ11 nm 

fðAu0Þ5:7KE5:ðNH2Þg C 87; dZ22:2 nmand radioactive fðAu0Þ5:8g C dZ22 nm 

f 198 ðAu 0 ÞKE5:ðNH2Þg C 87; dZ22:2 nm 

f 198 ðAu 0 Þg C dZ22 nm 

fðAu0Þ5:7KE5:ðNH2ÞC 209; dZ29:7 nmgand fðAu0Þ5:7KE5:ðNH2Þg C f 

dZ29:7 nm 

198ðAu0ÞKE5:ðNH2Þg C 209; dZ29:7 nm f198ðAu0ÞKE5:ðNH2Þg C dZ29:7 nm 

PSD: Positive Surface Dendrimer; NSD: Neutral Surface Dendrimer; Ng: Negative; CND: Composite Nanodevice 

a Denotes radioactivity. 

b Numbers of primary particles have been calculated from volume, assuming full space-filling. 

Method 28C. The ðAu 0 Þ 60:09KðPAMAM_ððE5ÞðE3:ðCOOHÞ 296ÞÞ K 

dZ11 nm 

, for short: 

fAuð0Þg K dZ11 nm, i.e., the dZ11 nm carboxylate surface (negatively charged at pHZ7.4) gold/dendrimer 

nanocomposite was synthesized from PAMAM_E5(E3.COOH) 9, a succinamic acidterminated 

tecto-dendrimer (see Table 28.1). Briefly, methanol solution of HAuCl 4 was added 

drop-wise to an aqueous solution of PAMAM_E5(E3.COOH) 9 tecto-dendrimer (MnZ120,000, 

functionalityZ296), resulting in a red solution of the corresponding dendrimer nanocomposite 

(Au/DZ60.09; composite gold contentZ8.94%). 

Method 28D. Positively charged dO10 nm dendrimer nanocomposites were made from generation 

three and generation five templated nanocomposites by simultaneously exposing them to 

gamma and neutron radiation in a nuclear reactor above Tg, the glass transition temperature of 


566 

10 nm 


FIGURE 28.9 TEM of single {(Au(0)nKE5 NH2} particles (marked with arrows) and various nanocomposite 

aggregates formed during sample preparation. (Reproduced from Balogh, L., Bielinska, A., Eichman, J. D., 

Valluzzi, R., Lee, I., Baker, J. R., Lawrence, T. S., and Khan, M. K., Chim. Oggi (Chem. Today), 20 (5), 35–40, 


the polymers, (for PAMAMs TgZ20–35 C). Under these conditions, the neutrons activate the 

trapped gold metal clusters in the {(Au 0 ) 5.76} primary composite nanoparticles, and the gammaradiation 

cross links the organic network. Irradiation times were varied from 30 min to 3 h. Level of 

activation of 197 Au within the allowable degree of polymerization of the dendrimer host was 

controlled by setting the neutron irradiation time and limiting the gamma dose by led filters. 

Gold ( 197 Au) captures neutrons very efficiently because of its large cross-section. 198 Au (t1/2Z 

2.69 days) decays predominantly by beta-radiation (99%, 0.96 MeV) with a small gamma component 

(0.98%, 1.1 MeV) into a stable 198 Hg isotope. The build-up of activity for any given isotope 

is a function of several intrinsic factors, including isotopic abundance and neutron cross-section as 

well as a series of irradiation parameters, including neutron flux and energy spectrum. Therefore, 

although the general principles are valid for all the reactors, detailed experimental parameters of the 

procedure have to be independently worked out. Under these experimental conditions, efficiency 

was 50% of the target (theoretical) value (Figure 28.10). 

There were no changes observed in nanocomposite particle size at low doses. At higher 

irradiation doses, the size of nanoparticles increased and the particle size distribution became 

somewhat wider. 

The exact mechanism of cross linking is yet unknown. It is speculated that radical mechanisms 

that are responsible for beta-alanine cross linking (as well as local heat effects that are known to 

result in retro-Michael synthesis) may result in the rearrangement of the molecular structure. 80 

Data indicate that the product (similarly to silver/dendrimer nanocomposites) 76 is composed of 

primary composite nanoparticles and their aggregates (Table 28.4). Volume weighted averages are 

directly proportional to %w/w concentrations. Nevertheless, interaction with biologic objects (e.g., 

cells) is determined by the number of interacting objects, thereby necessitating determination of 

number averages as well. By number averages, each sample is predominantly composed of single 

particles (Figure 28.11). 

The size of the cross linked primary particles depends on the irradiation conditions; brief 

exposure to neutrons and gamma rays does not change particle size (Figure 28.12, left side), but 

it results in low activities. Longer irradiation activates the gold and simultaneously melts the 



Measured activity microCi 

14 

12 

10 

8 

6 

4 

2 

Measd activity microCi 

Targeted activity microCi 

0 

0 5 10 15 20 25 30 35 

Irradiation time, min 

FIGURE 28.10 Comparison of targeted and measured specific activities of 198 Au as a function of irradiation 

time. (Reproduced from Balogh, L. P., Nigavekar, S. S., Cook, A. C., Minc, L., and Khan, M. K., PharmaChem 

2(4), 94–99, 2003. With permission.) 

dendrimers and cross links the primary particles into spherical and still soluble radioactive nanocomposites. 

Figure 28.12 shows the empirical correlation between particle diameter and irradiation 

time when pre-made {(Au 0 )6.53-PAMAM_E5.NH2} nanocomposites were treated in a reactor. 

In TEM, large particle structures that are probably responsible for the observed intense light 

scattering have been identified (Figure 28.13b). Visualization of these loose aggregates offers a few 

mechanistic clues. First, size and size distribution are unaffected when samples are irradiated with 

low doses in solution (Figure 28.13a). At higher doses, larger {Au} spherical particles form with a 

few superclusters present that are composed of primary {Au}n clusters loosely attached to each 

other by their surface. 

For approximate calculations, complete space-filling (100% packing) values for the cross 

linked primary particles and partial (50% packing) values for their aggregates were assumed. 

Based on volume averaged size data by DLS, a d Z 22.2G1.9 nm polymerized particle 

TABLE 28.4 

Particle Size Distributions of Soluble { 198 Au} C Nanocomposites After Irradiation of 

{(Au 0 )6.53-E5.NH2} (NICOMP Data) 

Sample {(Au 0 )6.53 

-E5.NH2}. A, B, C 

Mean 

Diam. 

(Nm) 

Volume Weighted Number Weighted 

Std. Dev. 

(Nm) 

Present 

(%) 

Mean 

Diam. 

(Nm) 

Std. Dev. 

(Nm) 

Present 

(%) 

Irradiation 

Time 

(Hr.) 

Zeta 

Potential 

(mV) 

fðAu0 C23:3 mV 

Þ6:53gnZ; dZ14:0G1:9 nm 14.0 1.9 99.4 13.4 1.4 100 1.0 C23.3 

Aggregate 453 45.1 0.6 0 0 0 


Þ6:53gnZ; dZ16:6G1:7 nm 16.6 1.7 94.6 15.8 1.9 99.9 1.5 C22.7 

Aggregate 73.6 10.3 5.4 70.4 6.2 0.1 


Þ6:53gnZ; dZ21:1G2:9 nm 21.1 2.9 91.2 20.2 2.2 99.9 2.0 C25.2 

Aggregate 111.2 17.2 8.8 104.9 4.0 0.1 

Data were measured after loss of radioactivity. 


568 

Rel % 

60 

50 

40 

30 

20 

10 

LBCB-6-2 Normalized 

0 

10 28 81 

Diameter (nm) 

Num wt % 

Vol wt % 

FIGURE 28.11 Comparison of number weighted and volume weighted distributions reveals that although this 

nanocomposite solution contains a few clusters, each sample is predominantly composed of particles of similar 

size. 

fðAu 0 Þ C 5:76 g 87; dZ22:2 nmcontains 87 primary (dZ5 nm) DNC units and 499 Au 0 atoms per particle. 

A dZ29.7G2.9 nm polymerized gold/dendrimer nanocomposite particle ðfðAu 0 Þ C 5:69 g 209; dZ29:7 nmÞ 

contains 209 primary (dZ5 nm) DNC units and 1123 Au atoms. The aggregate of this secondary 

nanoparticles (with a packing fraction of 50%) would be composed of dZ22 nm particles (each 

containing 489 Au atoms in this example) that would suggest the presence of 51.4 {( 198 Au)n} 

Diameter. nm 

25 

20 

15 

10 

5 

0 

60 

50 

40 

30 

20 

10 

0 0.5 1 1.5 2 2.5 

Time of irradiation, h 

0 

FIGURE 28.12 Average diameter of Peak 1 in aqueous solutions of gold nanocomposites as a function of 

irradiation time (in equivalent position). Extensively irradiated nanocomposite samples became insoluble in 

water. 



0 

25 

20 

15 

10 

5 

Num wt%


TABLE 28.5 

Particle Diameters of Radiation Polymerized Gold/Dendrimer Nanocomposites 

(Au/DZ5.69) According to Dynamic Light Scattering Data with NICOMP Data Analysis 

Sample 

Volume Weighted Number Weighted No. of Au atoms 

Mean 

Diameter 

(nm) Percent (%) 

Mean 

Diameter 

(nm) Percent (%) Per Host 

composite nanoparticles per aggregate (on average), containing 31,002 gold atoms in total. Data are 

shown in Table 28.5 and Table 28.6. The resulting nanocomposites preserved a net positive charge 

although measured zeta potential values (Table 28.6) were approximately half of what was 

measured for {(Au)56.61-PAMAM_E5.NH2} nanocomposites. 65 (The surface zeta potential for 

the latter nanocomposites was 41 mV as calibrated by dZ50 nm poly(styrene)sulfonate standards 

from Duke Scientific). Decrease of the surface positive charge indicates involvement of primary 

amine terminal groups in the polymerization reactions followed by their subsequent partial elimination 

from the host during radiation. The net positive zeta potential also supports the existence of 

the primary amine surface functions, explaining the formation of superclusters through 

surface attachment. 

28.1.5.7 Gold in the Nanocomposite Particles After Radiation Polymerization 

It is well known that the maximum of the surface plasmon resonances of noble metal nanocrystals 

in the UV–vis spectrum shifts toward longer wavelengths (from red to blue, i.e., toward lower 

energies) when particles aggregate. 81,82 However, the maxima in the UV–visible spectra of the 

{Au} nanocomposites are practically the same before and after radiation polymerization. The lack 

of shift indicates that the structural change is restricted to the host and the Au nanodomains remain 

isolated (Figure 28.14). 

28.1.5.8 Visualization of Gold/PAMAM Nanocomposite Uptake in Tumor Cells 

Gold mass 

Fraction (%) 

Primary particle: {(Au 0 )5.76}87 22.2 97.9 21.7 99.9 504 99.3 

Aggregate ({(Au 0 ) 5.76} 87) 12.5 51.5 2.1 51.4 0.1 3113 0.7 

In this demonstrative experiment, amine surface gold PAMAM nanocomposites were used to 

visualize {Au(0)}C in tumor tissue by TEM. One million B16F10 melanoma cells were 

TABLE 28.6 

NICOMP Size Distribution Data After Irradiation and Radioactive Decay for the Nanocomposite 

(Au/DZ5.76) Used in the Preliminary Therapeutic Experiments 

Sample 

Volume Weighted Number Weighted No. of Au atoms 

Mean 

Diameter 

(nm) 

Percent 

(%) 

Mean 

Diameter 

(nm) 

Percent 

(%) Per particle 

Mass 

Fraction (%) 

Primary particle: {(Au 0 )5.69}209 29.7 94.1 27.7 99.9 1123 97.5 

Aggregate ({(Au 0 )5.69}209)51.4 110.4 5.9 106.4 0.1 31,002 2.5 


570 

25 nm 

(a) (b) 

20.51 nm 

27.82 nm 

20.16 nm 

205.47 nm 

FIGURE 28.13 (a) The TEM image illustrates the unperturbed size distribution of the {Au5.7- 

PAMAM_E5.NH2} gold nanocomposite when irradiated with low doses. (b) Almost identical dZ22 polymerized 

{Au}n particles forming a loose ({Au}n)m aggregate. (Reproduced from Balogh, L. P., Nigavekar, S. 

S., Cook, A. C., Minc, L., and Khan, M. K., PharmaChem 2(4), 94–99, 2003. With permission.) 

subcutaneously injected on the dorsal surface of a C57Bl6/J mouse and tumor allowed to develop 

for two weeks. The mice were then injected with the nanocomposite solution. Tumor and other 

organs (liver and lung tissue) were isolated and ultra-thin sections examined on a Philips CM100 

electron microscope. Figure 28.15 shows a TEM image of one of these sections with three tumor 

cells. The image indicates that 1.25 h after the direct injection only a few larger gold nanocomposite 

clusters are present in the interstitium. Most of the smaller clusters were found in the nucleus and 

cytoplasm. During the internalization of gold nanocomposites, diffusion through the cell wall 

Extinction 

2 

1.5 

1 

0.5 

0 

200 

300 400 

500 600 700 

Wavelength, nm 


26. nm 

{Au(0)} d=22 nm 

{Au(0)} d=11 nm 

{Au(0)} d=5 nm 

800 900 1000 

FIGURE 28.14 Comparison of UV–visible spectra of dZ5 nm (dashed line), dZ11 nm (dotted line), and 

dZ22 (continuous line) nm {Au(0)} gold nanocomposites. 



FIGURE 28.15 Electron microscope image of mouse tumor tissue 1.25 h after {Au} injection. Unstained 

specimen, black dots represent gold nanocomposite clusters, dark areas within the nucleus are preferred for 

{Au}. (Reproduced from Balogh, L., Bielinska, A., Eichman, J. D., Valluzzi, R., Lee, I., Baker, J. R., 

Lawrence, T. S., and Khan, M. K., Chim. Oggi (Chem. Today) 20(5), 35–40, 2002. With permission.) 

appears to be the dominant mechanism although transfer through a nuclear pore has also 

been observed. 

28.2 BIOLOGY 

As previously shown, nanocomposite dimensions are in the size range of small cellular machinery. 

The interactions that devices this small will have with biologic systems is truly unknown, and 

mathematical or other guiding principles to determine how exactly these small devices will behave 

in complex biologic systems that have potential interactions at the intracellular, cellular, multicellular, 

organ, and multi-organ system level are yet unknown. Research has demonstrated that until 

understanding of this process is had, it will be critical to empirically test these interactions as simple 

assumptions about nanodevice behavior in biologic systems are currently often incorrect. 

Interactions of nanoparticles with biologic objects should typically be surface interactions. In 

other words, they depend on the surface characteristics of both the biologic entity and the nanoparticle. 

This has been borne out in several of this team’s experiments and will be detailed below. 

The nanoparticle’s size is important, and this may be due to cellular or macromolecular interactions 

that are affected by size such as the ability to move through membranes, pores, membrane junctions, 

or recognition by certain cells (such as macrophages) for uptake. Charge of the nanoparticle is 

important because the Coulomb interactions are the strongest physical forces between any two 

objects. Other three-dimensional changes within the nanodevices that can affect their surface 

size/charge would also then be expected to affect biologic interactions. 

Biodistribution, or how these nanodevices move throughout complex biologic multi-organ 

systems, is a dynamic property and must be looked at as a function of time. Biodistribution can 

be affected by many factors, including site of injection or uptake, blood flow/rate, uptake resident 

time in the blood and various organs and organ systems, and clearance from the body (urine, feces, 

and sweat, for example) to name only a few of these factors. Nanodevices intravascularly injected 

will travel first to the heart then lungs, then back to the heart, then throughout the rest of the body. 

The heart and lung have first exposure to the nanodevices. Certain organ systems may have 


572 

architecture (fenestrated structures for example) that may hasten or slow the time blood is exposed 

to the organ, the ability to enter organs, and the ability to stay in certain organs. Size/charge 

characteristics (particularly size) could have large influences on the route of clearance of these 

nanodevices. Detailed empiric biodistribution studies may eventually permit more detailed mathematical 

understanding of how different size/charge nanodevices will distribute over time in 

complex biologic systems. 

The vascular system is a critical organ system that needs detailed examination as it is the first 

organ system seen by intravascularly injected nanodevices and the main route of distribution to the 

rest of the body. Of particular importance to the study of cancer (and potentially other diseases such 

as macular degeneration, rheumatoid arthritis, etc.) is the angiogenic microvasculature. When these 

pathologic conditions occur or when a wound is healing, the normal existing microvasculature is 

stimulated to build a new microvasculature via budding off of the existing microvasculature. This 

process of neovascularization is called angiogenesis. It has potentially important implications for 

potential uses of nanodevices in pathologic and some non-pathologic biologic conditions as will be 

discussed below. A detailed understanding of the in vivo biodistribution of nanodevices will 

eventually permit the design of nanodevices that can specifically or selectively target specific 

organs or tissues in order to improve medical therapy. 

28.2.1 TARGETING AND NANODEVICE DELIVERY 


Targeting of tumor cells may be performed by either selective or specific ways as described in 

related literature. Direct receptor specific targeting is the one most often thought about, and 

nanocomposites permit the optimization of surface properties (via modulation of charge and attachment 

of targeting moieties) to maximally permit this direct targeting. Targeting groups can be 

conjugated to the host dendrimer’s surface 27 to allow the imaging agent to selectively bind to 

specific sites such as receptors on tumor cells to improve detection. For example, there are a 

number of cell surface receptors that are known to be over expressed on cancer cells, including 

epidermal growth factor receptor (EGFR) and the laminin receptor VLA-683. Nanocomposites 

could be constructed to specifically target these receptors. 

Nanodevices can also be constructed to target exposed receptors on the angiogenic microvasculature. 

Over the last few years, data have emerged supporting the concept that the 

microvasculatures of tissues and/or organs also have antigenic surface differences. Using a 

phage display technique, investigators have isolated peptide fragments that specifically bind to 

different tissue or organ microvasculatures in mice. 84–86 As part of this work, they also isolated 

peptides that specifically bind to angiogenic tumor microvasculature. 87–89 One of these angiogenic 

microvascular targeting peptides isolated is a three amino acid peptide, arginine–glycine–aspartate 

(RGD). RGD has been used to target chemotherapy and other agents to the angiogenic tumor 

microvasculature. 87,90,91 Radio-labeled linear or cyclic forms of the RGD peptide itself have 

been used to target the tumor microvasculature, permitting imaging with PET or SPECT. 92–97 

One could envision the use of RGD or other small molecules binding angiogenic microvasculature 

and present on the surface of the composite nanodevices to directly target the microvasculature. 

These studies are currently underway in this team’s laboratories. 

Non-target specific or non-selective targeting is another way to affect biodistribution of the 

nanodevices by not relying on the binding to a specific receptor but by taking advantage of the 

existing structural properties of complex biologic systems. It is well known that the angiogenic 

microvasculature is very different from the fully formed microvasculature. It is leakier than normal 

microvasculature, and there is also evidence that the tumor angiogenic microvasculature often lacks 

pericytes and has aberrant morphologies. 98–101 These differences between the tumor microvasculature 

and the normal microvasculature may result in enhanced permeability and retention effect 

(EPR), suggested first by Maeda. 102 The accumulation of macromolecules in the tumor was also 

found after i.v. injection of an albumin-dye complex (MwZ69,000) as well as after injection into 



normal and tumor tissues. The complex was retained only by tumor tissue for prolonged periods. 

With liposomes, the most effective size preferably retained in the tumor was suggested to be around 

50 nm. The Duncan group has also shown 103 that appropriately sized dendrimers and dendrimercis-platin 

complexes may become trapped in tumor vasculature as a result of the EPR effect. 104 

Using PAMAM dendrimers, it has also been demonstrated that polymers of different sizes have 

different permeability or cellular uptake, depending on the size (generation) and surface of the 

dendrimer (discussed in more detail below). For future therapy, these size and charge characteristics 

are exploited to specifically deliver nanoparticles through the tumor microvasculature 

and to the tumor or directly to the microvasculature itself using additional surface recognition to 

further improve avidity. 

28.2.1.1 Quantitative In Vivo Biodistribution of Dendrimer and Gold Composite 

Nanodevices (CND) 

As previously explained, composite nanodevices have several potential advantages over current 

technologies for the delivery or radiation or other therapies to tumors or the tumor microvasculature. 

In particular, even the smallest gold composite nanodevice can deliver greater than ten gold 

atoms per device to a target. This is a log fold more radiation than previously possible with the usual 

radioactive antibody techniques and offers a real potential to have greater than a log fold more 

radiation delivery per device than previously possible. The composite nanodevice architecture 

permits the simultaneous optimization of the carried inorganic component (gold or other inorganic 

material) and the dendrimer network (with or without receptor-specific targeting), providing the 

surface components of the fully formed nanodevice. To think of this clinical point of view, for 

example, targeting the nanodevice by the covalent linking of specific moieties to its surface and 

then carrying radioactive gold for imaging/therapy applications, the gold placement would not 

interfere with the targeter functioning on the surface. This optimization of the different compartments 

is a critical advantage of the composite nanodevices compared to other structures. 

28.2.1.2 Related Techniques in the Literature 

Illustrative of a few of these points is a study by Haifield et al. in the early 1990s. 55 This group 

carried out a partially successful attempt to increase the gold carrying content of antibodies by 

covalently attaching gold clusters using sulfhydryl linkages to antibody fragments, and it then 

carried out biodistribution studies of these devices in mice. They were able to get some of the 

gold clusters averaging about 11 gold atoms to bind to various antibody fragments. The characterization 

of these clusters was done with TEM, and much more detailed characterization using several 

other techniques would be needed to be certain of the actual complexes formed. The size of the final 

assembly was not presented. They were able to make an antibody fragment-Au-cluster device that 

retained about 80% of its original binding activity (because of the interference of the gold clusters 

with the antibody fragment binding site). To stabilize the clusters, 21 phenyl groups were used, and 

this hydrophobic layer around the Au clusters caused increased uptake arguably in the blood cells. 

There was also very high uptake (trapping of clusters) in the kidney, probably as a result of the size 

of the antibody-Au cluster constructs. In both cases, the Au cluster chemistry/architecture interfered 

with the targeting component of the devices. More recent work on colloidal gold preparations by 

Paciotti et al. 56 examined the in vivo biodistribution of a colloidal gold nanoparticle devices made 

of thiol–PEG (polyethylene glycol) and gold and recombinant human TNF (PT–cAu–TNF). The 

mechanism of the TNF linkage to the colloidal gold could not be delineated, but the preparations 

did appear to be 33 nm in size, and crude biodistribution measurements in a few organs showed 

relatively rapid clearance from the blood and deposition of the TNF in tumors. Much of the 

biodistribution of the nanoparticle was done by observing the color of organs as they turned 

more purple as they picked up the nanoparticle (possibly because of the formation of large gold 


574 

nanoparticle aggregates that shifts the maximum of the plasmon peak toward blue). The data were 

not quantitative for whole particle determinations. The measurement of TNF in a few organs was 

quantitative. This nanoparticle appeared to have some interesting data supporting improved 

delivery to tumors compared to the non-PEGylated form of the nanodevice. It is unclear from 

the work, however, how this approach could be extended to permit other targeting with different 

agents or what improvements need to be made. There was no indication of how the Au–PEG–TNF 

components could be modified or optimized. Given the lack of knowledge of the actual chemistry 

holding the device together for the short few-hour time course of the experiment, it is unclear if this 

system can be modified or improved. These reports emphasize the advantages possible with the 

composite nanodevices, the need for much more detailed chemical characterization and biological 

understanding of how these devices distribute throughout the body, and how simple modifications 

may affect this biodistribution. 56,105 

Quantitative and detailed in vivo biodistribution data for dendrimer nanoparticles and, for the 

first time, composite nanodevices in tumor mouse model systems are presented below. Although 

there are many theories on how nanodevices might interact with complex animal systems, this team 

decided to empirically measure nanodevice biodistribution with several planned modifications of 

the nanodevices as a way to eventually begin to understand guiding principles of how nanodevices 

and, in particular, nanocomposites interact with organs/tissues. This data will eventually be used to 

develop a more detailed mathematical understanding of the process and to aid in the design of future 

nanodevices aimed at specific biologic understanding or therapy. 

As part of this approach, these groups carried out the first detailed and quantitative biodistribution 

studies on dendrimer nanodevices in tumor model systems. 77 They then expanded these studies 

to examine the first detailed and quantitative biodistribution studies of nanocomposites. 4 Important 

initial principles of the interactions of nanocomposites were determined, and predictions were made 

about how to examine nanodevices in general in complex biologic systems. 4,77,79 

28.2.1.3 The Instrumental Neutron Activation Analysis (INAA) Method 


Gold content of biologic samples by Au content of the nanodevices and tissue samples was 

determined by direct neutron irradiation. This approach was possible because (a) in the absence 

of oxygen, the gold/dendrimer nanocomposites are stable for a number of months (metal nanoclusters 

are effective oxygen-transfer catalysts), (b) PAMAMs do not break down in vivo because of 

enzymatic activity, and (c) the natural abundance of gold in mice and humans is close to zero. 

In the INAA technique, samples are activated by irradiation with neutrons, usually within the 

high flux fields of a nuclear reactor. Specific isotopes become radioactive by neutron capture-type 

nuclear reactions. After removal from the reactor, samples are allowed to decay (cool) to permit 

unwanted short-lived activity (for example, from Na) to diminish. Gamma-ray spectra from activated 

samples are then measured using solid-state detectors. 24 Specific isotopes may be identified 

by their gamma rays of characteristic energy, and quantitative determinations are made by 

comparing peak areas in the spectrum to those of a standard reference material that has been 

irradiated and counted under identical conditions. 

Advantages of this analysis method are that (a) determination of CND biodistribution is very 

precise, (b) the biologic experiments can be performed on their own timescale and schedule, 

(samples can be prepared, accumulated, stored, and analyzed upon reactor availability), (c) high 

sensitivity may be achieved for gold, and (d) there is no biohazard material present after irradiation. 

However, it is a disadvantage that the dried and irradiated tissues are not suited for further study 

(TEM, pathology, immunology, etc.) The build-up of activity for a given isotope is a function of 

several intrinsic factors, including isotopic abundance and neutron cross-section as well as a series 

of irradiation parameters, including neutron flux and energy spectrum. Because irradiation parameters 

can vary slightly over time within a range of operating conditions, quantification of the 

activated isotope is rarely approached through direct calculation. Rather, elemental concentrations 



are determined through comparison on a weight-ratio basis from activities (measured as gamma 

counts-per-second) generated in a standard reference material of known composition, assuming all 

other parameters are held constant, including irradiation time, flux, decay time, counting time, and 

detector geometry. 

28.2.1.4 Determination of Gold Content Using INAA—Precision and Sensitivity 

All tissue samples were lyophilized and analyzed either at the Ford Nuclear Reactor of the Phoenix 

Memorial Laboratory (University of Michigan) or at the Oregon State University Radiation Center 

(OSU-RC, Corvallis, OR). A series of six gel check standards prepared with known gold concentrations 

(in 7% gelatin to mimic the organic tissue samples) were prepared to monitor measurement 

precision and sensitivity. Four replicates of a gold standard (NBS-SRM-2128-1, 1% Au by weight) 

prepared by pipetting approximately 100 mg of the liquid standard on to filter paper placed in 

sample vials were similarly weighed and desiccated. Samples and standards were irradiated for 2 h 

in core-face location 112B with a nominal thermal neutron flux of 10EC11. Following an initial 

2-day decay, a 15,000 s (live-time) count of gamma activity for Au-198 was recorded for each 

sample using a HPGe detector (33% relative efficiency) with a 4 00 counting geometry. Gold 

concentrations were determined through direct comparison on a weight ratio basis with the mean 

activity generated in four replicates of the Au standard. Sensitivities of !75 ppb were obtained, 

corresponding to 10 ng of Au (Figure 28.16). This method permits detailed and quantitative biodistribution 

measurements in mouse tumor model systems. 

28.2.1.5 Biodistribution of Positive and Neutral Surface Generation Five 

Dendrimers—Non-Specific Uptake 

In the first study of PAMAM dendrimers, the surface of PAMAM dendrimers were labeled with 

tritium 77 and studied the biodistribution of PAMAM_E5:ðNH 2Þ 44ðNHCOCH 3 Þ 66 a tritiated generation 

five (5 nm) positive surface dendrimer and PAMAM_E5:ðNHCOCH 3 Þ 110 neutral surface 

PAMAM dendrimers (PSD and NSD, respectively) in a mouse melanoma tumor model and in a 

prostate cancer mouse model. The feasibility of dendrimer delivery and subsequent quantification 

was also tested. Systemic administration of NSDs and PSDs via tail vein injection were carried out 

in tumor bearing mice. At different time points, the mice were euthanized and their organs 

Calculated Au, (ng) 

10 6 

10 5 

10 4 

1000 

100 

10 

1 

0.1 

0.1 

y = 0.26864 * x ∧ (1.1488) R= 0.99971 

1 10 100 1000 10 4 

Measured Au, (ng) 

FIGURE 28.16 Comparison of Au measured by INAA and the theoretical (known) Au content measured in 

gel sample (gold concentrations were determined through direct comparison on a weight ratio basis with the 

mean activity generated in four replicates of the Au standard). 


10 5 

10 6

576 

harvested, weighed, and processed. The in vivo distribution of dendrimers was then determined via 

liquid scintillation counting of recovered tritium from the organs. 

Although neither particles localized selectively to tumor tissue, they did distribute to all organs 

tested and were recoverable within one hour post-injection. The vascular delivery of dendrimers 

could be homogenous in the organ (as seen in multiple samples of the liver analyzed). Both dendrimers 

rapidly cleared from blood, the % ID/g organ levels for PSD and NSD being 20.6 and 21.6 at 

5 min, 14.3 and 9.9 at 1 h, 1.7 and 0.9 at 1 day, 0.3 and 0.1 at 4 days, and 0.2 and 0.1 at 7 days postinjection, 

respectively. These organic nanoparticles localized to all major organs and tumor tissue 

and were rapidly cleared from the blood. Following an initial rapid clearance during the first day postinjection, 

concentrations were maintained at a relatively stable level in every tissue with a very slow 

decline over time. Lungs, kidneys, liver, and heart exhibited the highest uptake within an hour. Levels 

in the spleen and tumor followed this, and the negligible amounts were found in the brain tissue. More 

detailed tables and graphs of both PSD and NSD biodistributions are available in this team’s previous 

publication. 77 The PSD dendrimer nanoparticle biodistribution is shown in Figure 28.17. NSD 

biodistribution is very similar, but with lesser amounts throughout. In all organs and tissues tested 

(including tumor) the ratio of PSD:NSD was higher than 2:1, most likely because of increased 

retention of the positively charged dendrimers relative to neutral derivatives. It shows, in a quantitative 

manner, a global and differential effect of charge on organ and tissue biodistribution when the 

dendrimers are almost the same size (dZ5 nm). 

In another experiment, the nanoparticle biodistribution was determined in another tumor model 

system (human DU145 prostate cancer xenograft), and comparison was made to the previously 

determined B16 melanoma tumor model system (see Figure 28.17). The results are very similar in 

both models, supporting the general validity of the findings to multiple tumor and mouse types. 

In detailed excretion studies, it was shown that the dendrimers were excreted via both urine and 

feces (Figure 28.18). A total of 48.3% of NSD and 29.6% of PSD were excreted via urine over 

seven days. Within the first two hours, urinary NSD excretion was more than three-fold that of PSD, 

but at later sampling points, the daily excretion of both dendrimers was equivalent. This correlated 

with the increased retention of PSD seen in organs and tissues. Dendrimers were excreted via feces 

to a much lesser extent with a total of 5.41% of NSD and 2.96% of PSD recovered over seven days 

50 

45 

40 

35 

30 

25 

20 

15 

10 

5 

0 

(a) Biodistribution of 5 nm PSD in C57BL6/J mice 

Blood 

Lungs 

Heart 

Liver 

Pancreas 

Time (h) 

Spleen 

Kidney 

Tumor 

1 

24 

96 

Brain 

8 

7 

6 

5 

4 

3 

2 

1 

0 

Blood 

(b) Biodistribution of 5 nm PSD in nude mice 

Lungs 

Heart 

Liver 

Pancreas 

Time (h) 

FIGURE 28.17 Organ biodistribution of positive surface 5 nm nanoparticle (PSD) depicted as percent 

injected dose recovered per gram of organ (% ID/g). Panel (a) shows biodistribution in C57BL6/J mice 

subcutaneously carrying B16 melanoma tumors on the dorsal surface. Panel (b) shows the same nanoparticle 

injected into nude (nu/nu) mice carrying the human prostate carcinoma line DU145. nZ6 for panel (a) and 

nZ3 for panel (b). 



Spleen 

Kidney 

Tumor 

1 

24 

96 

Brain


25 

3 

20 

15 

NSD 

2 

10 

5 

PSD 

1 

0 

Time (Hours) 2 8 24 48 72 96 120 144 168 

0 

Time (Hours) 

% Injected dose 

NSD 24.92 12.82 6.29 2.76 0.72 0.33 0.18 0.14 0.15 

PSD 7.84 10.84 6.24 2.81 1.00 0.37 0.22 0.15 0.15 

(a) (b) 

(Figure 28.18). The dendrimers were mainly excreted via urine with a significant amount of urinary 

excretion occurring within 24 h post-injection. This was the first quantitative demonstration of the 

manner in which dendrimers are excreted from the body (urine) and of how this occurs over time. 

28.2.1.6 Biodistribution of 5 nm Gold Nanocomposites Compared with Dendrimer 

Templates 

As the next step in this research, this team examined composite nanodevices of differing surface. 4 It 

compared the biodistribution of 5 nm composite nanodevices with those of their respective 

dendrimer templates of the same surface charges and size. It found that although charge affected 

biodistribution of the same size nanodevices, surprisingly, the gold/dendrimer composite nanodevices 

had a very different organ/tissue biodistribution pattern in mice compared to the dendrimer 

nanoparticles of similar size and surface charge. 

28.2.1.7 Biologic Differences Between Dendrimers and Gold Composite Nanodevices 

This comparison is most accurate and particularly striking in the case of the acetamide (neutral) 

surface nanodevices where the size and surface charge are the same. There is significant and 

consistent high-level accumulation over time (selectivity) in the liver and spleen of mice intravenously 

injected with the gold nanocomposites (Figure 28.19a). In contrast, no such organ specific 

accumulation (selectivity) was seen with the NSD (Figure 28.19b). 

The overall (global) effects of surface charge are also different when the dendrimer nanoparticles 

are compared with the CNDs. With the dendrimers (Figure 28.19a and Figure 28.20a), and as 

discussed above, the partially positive surface macromolecules distributed into all organs at higher 

levels (about 2-fold higher levels) than that seen with the fully acetylated (neutral surface) 

dendrimer nanoparticles. 77 There were no significant organ-to-organ differences (or selectivity) 

seen. With the gold composite nanodevices, however, there are large organ-to-organ differences 

(selectivity) seen (Figure 28.19a and Figure 28.20b). There is significant spleen and liver accumulation 

(selectivity) seen with the fðAu 0 Þ 9:08KPAMAM_E5:ðNHCOCH 3Þ 120g C dZ5 nmneutral surface 

nanodevices over a prolonged time period. With the positive surface fðAu 0 Þ 9:08KPAMAM_E5: 

ðNH 2Þ 120g C dZ5 nmnanocomposite, there was significant accumulation noted in the kidney. 

Finally, a further major difference noted was that the unmodified primary amine terminated 

PAMAM dendrimer was previously shown to be toxic in vivo (causing the death of mice) as 

discussed in the introduction, and this is exactly why partially acetylated PAMAMs were used 

for in vivo studies as PSD. With the gold/dendrimer composite nanodevices, however, unmodified 

hosts were used to safely prepare respective nanocomposites for these mouse tumor model studies. 

Although the exact reasons are yet unknown, dendrimers and dendrimer nanocomposites are 

% Injected dose 

NSD 

PSD 

2 8 24 48 72 96 120 144 168 

NSD 0.40 2.63 1.43 0.74 0.03 0.03 0.05 0.05 0.05 

PSD 0.76 0.86 0.41 0.42 0.1 0.04 0.05 0.06 0.26 

FIGURE 28.18 Excretion of positive surface dendrimers (PSD) and neutral surface dendrimers (NSD) via 

urine and feces (a and b, respectively) at time points indicated. The data obtained from B16F10 melanoma 

mouse model (nZ5) are presented as a percentage of injected dose excreted per mouse. 


578 

Biodistribution of 5 nm Positive Au-Composite 

Biodistribution of 5 nm Partially Positive Dendrimer 

Nanodevice in the B16 mouse melanoma model 


45 

45 

40 1 h 

40 1 h 

35 1 day 

35 1 day 

30 4 days 

30 4 days 

25 

25 

20 

20 

15 

15 

10 

10 

5 

5 

0 

BLD BRN PCS TMR HRT KDY LVR SPN LNG 

0 

BLD BRN PCS TMR HRT KDY LVR SPN LNG 

(a) (b) 

%ID/g organ 

FIGURE 28.19 Organ/tissue biodistribution of neutral surface Au-composite nanodevice or (Au0)9-PAMA- 

M_E5.(NHCOCH3)110} (a) and neutral surface dendrimer or PAMAM_E5.(NHCOCH3)110 (b) in tumor 

bearing mice. Abbreviations: BLD, blood; BRN, brain; PCS, pancreas, TMR, tumor; HRT, heart; KDY, 

kidney; LVR, liver; SPN, spleen; and LNG, lung. (The (b) has been reproduced from Nigavekar, S. S., 

Sung, L. Y., Llanes, M., El-Jawahri, A., Lawrence, T. S., Becker, C. W., Balogh, L., and Khan, M. K., 

Pharm. Res., 21(3), 476–483, 2004. With permission.) 

different enough to yield different biologic interactions (even though they are similar by charge and 

size in materials characterization), resulting in a different toxicity profile. 

The surface-to-volume ratio of nanodevices is much greater than for micron objects, enabling 

the surface to govern the interactions of nanoparticles with proteins, carbohydrates, or other macromolecules 

within the organs, tissues, or cells. It was previously believed that the surface of the 5 nm 

positive or NSDs would largely govern the biodistribution of the gold/dendrimer composite nanodevices. 

Even though gold nanocomposites of primary amine terminated PAMAMs typically form 

mixed structures, 41 the fact that the fðAu 0 Þ 9:08KPAMAM_E5:ðNHCOCH 3Þ 120g C dZ5 nm and 

fðAu 0 Þ 9:08KPAMAM_E5:ðNH 2Þ 120g C dZ5 nm nanodevices have very different biodistribution within 

the mouse tumor model system in comparison to the dendrimer templates indicates that, from a 

biologic point of view, these devices interact differently with macromolecules present on or in 

organs/tissues. A classic three-dimensional conformational change (as can be seen with proteins) 

and perhaps affecting these surface interactions cannot account for the difference in biodistribution 

%ID/g organ 

45 

40 

35 

30 

25 

20 

15 

10 

5 

0 

Biodistribution of 5 nm Neutral Au-Composite 


1 h 

1 day 

4 days 

%ID/g organ 

%ID/g organ 

45 

40 

35 

30 

25 

20 

15 

10 

5 

0 


Biodistribution of 5 nm Neutral Dendrimer 


1 h 

1 day 

4 days 

BLD BRN PCS TMR HRT KDY LVR SPN LNG BLD BRN PCS TMR HRT KDY LVR SPN LNG 

(a) (b) 

FIGURE 28.20 Organ/tissue biodistribution of positive surface Au-composite nanodevice or {(Au0)9- 

PAMAM_E5.(NH 2) 110 (a) and partially positive surface dendrimer or PAMAM_E5.(NH 2)4 4 (NHCOCH 3) 66 

(b) in tumor bearing mice. Abbreviations: BLD, blood; BRN, brain; PCS, pancreas; TMR, tumor; HRT, heart; 

KDY, kidney; LVR, liver; SPN, spleen; and LNG, lung. (The (b) has been reproduced from Nigavekar, S. S., 

Sung, L. Y., Llanes, M., El-Jawahri, A., Lawrence, T. S., Becker, C. W., Balogh, L., and Khan, M. K., Pharm. 

Res., 21(3), 476–483, 2004. With permission.) 



as all the studied particles are spherical. The increased rigidity and density of {Au(0)} composite 

nanodevices compared to the relative flexibility of the template dendrimers may explain some of 

this difference. Alternatively, perhaps another changed physiochemical property may partially be 

responsible for some differences in biodistribution. This aspect is presently under investigation. 

The results again indicate that those working on nanoparticles for biologic and medical uses 

need to understand in detail how nanodevices truly behave in vivo and that new nanodevices will 

have to be rigorously tested for biodistribution, toxicity, and other critical measures needed for the 

use of these devices in biologic systems. As future multifunctional nanodevices are made, it will be 

critical to carry out detailed biologic testing of these devices with respect to changes in charge, size, 

complexation, and surface substitution to gain an in depth understanding of how they will behave in 

these systems. Studies on dendrimers and composites have shown that this rigorous testing is 

critical and will be needed for other nanomaterials as well. 

28.2.1.8 The Importance of Size on Nanocomposite Biodistribution 

This team conducted a series of further studies that are currently in submission for publication 

where the effect of size and charge on biodistribution of nanocomposites is further examined. It 

examined 5, 11, and 22 nm nanocomposites with varying surface charge (partially positive, 

negative, and neutral). A key finding of these studies was that there are different organs that are 

selectively uptake charged nanoparticles, depending on the size and charge tested (see 

Figure 28.21). Importantly, when surface charge was kept constant and the size of the nanodevice 

was varied as with the 5 and 11 nm carboxylate (negative) surface nanocomposites, different 

nanodevice biodistribution was seen with differential organ selective uptake noted. 

Nanodevice levels in blood are higher for the fAuð0Þg K dZ5 nmgold nanocomposites than they are 

for the fAuð0Þg K dZ11 nm gold nanocomposites. Smaller size might preclude quick clearance from 

blood circulation, perhaps by uptake of the nanodevices into circulating blood cells. Nanodevice 

accumulation in the spleen appears to be determined by size as the fAuð0Þg K dZ11 nm negative surface 

gold nanocomposites show much higher levels of accumulation in the spleen than the 5 nm negative 

% ID/gorgan 

300 

250 

200 

150 

100 

50 

0 

5nm 

NgCND 

Biodistribution of 5 and 11 nm negative surface Au-composite 

nanodevices in the B16 mouse melanoma model 

11nm 

NgCND 

5nm 

NgCND 

11nm 

NgCND 

5nm 

NgCND 

11nm 

NgCND 

5nm 

NgCND 

11nm 

NgCND 

5nm 

NgCND 

Liver Spleen Lungs Kidneys Tumor 

Mouse organ 

FIGURE 28.21 Biodistribution of 5 and 11 nm negative surface Au-composite nanodevices in C57BL6/J 

mice bearing B16F10 melanoma. Shown here are nanodevice accumulation in tumor as well as in organs that 

show the highest accumulation of nanodevice that include the liver, spleen, lungs, and kidneys. Nanodevice 

levels were tested at times 5 min and hours 1, 24, and 96 post-injection. 


5 min 

1 h 

24 h 

96 h 

11nm 

NgCND

580 

surface Au-composite nanodevice. Liver accumulation was also consistently higher in the 11 nm 

nanocomposites. These studies indicate that size can affect distribution of the nanodevices with 

similar surface charges. This will have to be taken into consideration if this nanodevice were to be 

used to deliver radiation or toxic therapeutics, especially in the absence of a tumor targeting moiety 

on the surface of the nanodevice. This is the first set of data demonstrating the critical importance of 

the size of the nanodevices on their biodistribution in complex organisms. 

28.2.1.9 Biodistribution Summary 

Detailed quantitative biodistribution studies demonstrated the critical importance of surface size 

and charge on the interactions of nanocomposites with complex biologic systems. This team was 

able to empirically determine how differing size and charge affects this biodistribution, showing 

that selective organ uptake can be achieved with modulation of these parameters. It also showed 

that biodistribution profiles of Au-composite nanodevices were different from profiles of dendrimers 

of the same size and charge. These studies also revealed that an important principle of 

nanodevice design is to be fully aware of these parameters, including the importance of flexibility 

and density. Importantly, this team also demonstrated that it is difficult to predict what size/charge 

modification would produce a given biodistribution. The behaviors appear to be complex at this 

time. Completion of its ongoing research will provide a deeper insight into the biodistribution of 

nanoparticles of differing size and charge and create an important background to examine targeted 

devices. Without this detailed type of baseline study, there would be no way to ensure true targeting 

is occurring as envisioned targeting may be more a result of selective uptake because of size, 

charge, and other properties, and they are not due to the targeting moiety on the nanodevice. 

Understanding nanodevice behavior in complex animal systems will help to design the best nanodevices 

for targeting and suggest ways to protect organs from undesired effects. The detailed 

studies above will also allow researchers to begin a more detailed mathematical and pharmacokinetic 

analysis to explain nanodevice size/charge effects on biodistribution and permit them to make 

educated predictions as to their behavior in biologic systems in the future. 

28.2.2 TOXICITY STUDIES 

28.2.2.1 Long-Term Toxicity Studies of Tritiated Dendrimer Nanoparticles 


Initial long-term toxicity studies were carried out on both the dendrimer nanoparticles and on 

several of the composite nanodevices as indicated in Table 28.7. As in vitro toxicity data cannot 

yet be directly transposed to in vivo conditions, this team has carried out initial toxicity studies on 

the tritiated dendrimers. Because tumor-bearing mice would die as a result of their tumor-burden, 

long-term toxicity/biodistribution studies were carried out with normal C57BL/6J mice. PAMAM 

_E5:ðNH 2Þ 44ðNHCOCH 3 Þ 66 generation 5 PAMAM dendrimer derivative with a partially positive 

surface (PSD) and a neutral surface PAMAM_E5:ðNHCOCH 3 Þ 110 dendrimer (NSD) were evaluated 

for in vivo toxicity in 6–8 week old male C57BL6/J mice. The mice were observed for a period 

of 75 days. There were five mice in each group with a control group that was given PBS. Clinical 

toxicity tables were kept on each set of mice, and daily notations were made of any signs of 

weakness, lethargy, dehydration, change in mobility, or reflex. Weights of mice-given dendrimers 

did not differ from weights of mice in the control group (see Figure 28.22), and none of the mice 

died during observation. The labeled dendrimers were detectable in all organs tested (see Table II in 

Nigavekar et al. 2004). 77 This demonstrated that the nanoparticles could distribute to all tested 

organs and remain there at relatively stable levels for at least 12 weeks. The mice appeared healthy 

at that time as well, and no long-term clinical toxicity or weight loss was observed. The mice 

weights increased over the first 15 days and then were stably maintained with slight increment over 

the period observed. This team’s conclusion was that these dendrimers are non-toxic in this 

animal model. 



TABLE 28.7 

Male C57BL6/J, 6–8-Week-Old Mice Were Injected with Tritiated Dendrimer Nanoparticles 

(5 nm PSD and NSD) and Au-Composite Nanodevices (5 nm PCND, 11 nm NgCND, and 

22 nm PCND) and The Were Monitored Over a Period of Time for Change in Weights 

Toxicity analysis 

Analysis of Mice Weights 5 (nm) PSD 5 nm NSD 5 nm PCND 

11 nm 

NgCND 22 nm PCND 

Time course of analysis 

(in days) 

Analysis of clinical 

75 75 75 49 49 

Toxicity O O O O O 

The mice were also monitored for signs of clinical toxicity such as weakness, lethargy, dehydration, change in mobility, or 

reflex. Death, as an end-point was noted as well. 

Weight (g) 

40 

30 

20 

10 

Weight (g) 

40 

30 

20 

10 

Weights of C57BL6/J 

mice given PSD 

0 

1 3 6 11 19 34 53 71 

Time (days) 


mice given PBS 

0 

1 4 11 27 53 75 

Time (days) 

PSD 1 

PSD 2 

PSD 3 

PSD 4 

Weight (g) 

40 

30 

20 

10 

Control 1 

Control 2 

Control 3 

Control 4 

Control 5 


mice given NSD 

0 

1 3 6 11 19 34 53 71 

Time (days) 

NSD 1 

NSD 2 

NSD 3 

NSD 4 

NSD 5 

FIGURE 28.22 Long-term monitoring of weights of non-tumor bearing C57BL/6J mice given (a) PBS as 

control (nZ5), (b) PSD (nZ4) and (c) NSD (nZ5). Individual weights of each mouse have been plotted as a 

function of time (days). PBS, phosphate buffered saline; PSD, positive surface dendrimer; NSD, neutral 

surface dendrimer. The notation PSD 1, for example, means the positive surface dendrimer injected into 

mouse 1. 


582 

28.2.2.2 Long-Term Toxicity of {Au(0)} Gold Composite Nanodevices 

Very limited data are available on the toxicity of noble metal dendrimer nanocomposites. Some 

preliminary data in cell culture systems provide some understanding of metal composite nanodevices 

in vitro. 5,76 It appears that toxicity is very similar to that of the dendrimer template, i.e., if the 

template is non-toxic, the noble metal composite nanoparticles are also non-toxic. 76 

To gain the insight into the in vivo toxicity of composite nanodevices, the team completed 

detailed in vivo toxicity analysis of fAuð0Þg C dZ5 nmthe 5 nm positively charged CND, fAuð0Þg K dZ11 nm 

the 11 nm negatively charged CND, and fðAu 0 Þ 5:8g C dZ22 nm the 22 nm positively charged CND in 

healthy mice. The mice were injected with the nanodevices as described above for the tritiated 

dendrimers. In brief, the 5 nm PCND, the 11 nm NgCND, and the 22 nm PCND were all evaluated 

for toxicity in 6–8 week old male C57BL6/J mice. Mice given the 5 nm nanodevice were observed 

for 75 days, whereas those given the 11 nm and 22 nm devices were observed for 49 days. There 

were five mice in each group with a control group that was given PBS. Daily toxicity tables were 

kept on each group of mice, and any signs of weakness, lethargy, dehydration, change in mobility, 

or reflexes noted. There was no clinical toxicity found. Weights of mice that were given dendrimers 

did not differ from weights of mice in the control group (see Figure 28.23), and none of the mice 

died during observation. 

28.2.2.3 Treatment of Tumors in Mouse Model Systems 

Reactive encapsulation permits the entrapment of different radionuclides within composite nanoparticles 

of defined size and targeting properties. An advantage of this approach would be the 

development of an image guided therapy where the radioactivity delivered to a tumor can be 

increased either by increasing the particle size or by increasing the number or specific activity of 

the guest atoms without destroying the targeting ability of the nanocomposite device. Additionally, 

Weight (g) 

Weight (g) 


mice givenPBS 

30 

25 

20 

15 

10 

5 

0 

1 3 7 12 18 36 40 46 

Time (days) 

40 

30 

20 

10 

0 

Weights of mice given 

5nm Au-PCND 

1 5 7 14 28 43 58 75 

Time (days) 

Control 1 

Control 2 

Control 3 

Control 4 

Control 5 

5nm PNCD 1 

5nm PNCD 2 

5nm PNCD 3 

5nm PNCD 4 

5nm PNCD 5 

Weight (g) 

Weight (g) 

35 

30 

25 

20 

15 

10 

5 

0 

30 

25 

20 

15 

10 

5 

0 



11nm Au-NgCND 

1 3 7 12 18 36 40 46 

Time (days) 


22nm Au-PCND 

1 3 7 12 18 36 40 46 

Time (days) 

NgNCD 1 

NgNCD 2 

NgNCD 3 

NgNCD 4 

NgNCD 5 

22nm PNCD 1 

22nm PNCD 2 

22nm PNCD 3 

22nm PNCD 4 

FIGURE 28.23 Long-term monitoring of weights of non-tumor bearing C57BL/6J mice given PBS (nZ5) as 

control (nZ5), 5 nm Au-PCND (nZ5), 22 nm Au-PCND (nZ4), and 11 nm Au-NgCND (nZ5). Individual 

weights of each mouse have been plotted as a function of time (days). Au-PCNDZgold positive surface 

composite nanodevice; Au-NgCNDZ gold negative surface composite nanodevice; PBSZ phosphate 

buffered saline. The notation PCND 1, for example, means that particular nanodevice injected into mouse 1. 



varying the particular nature of the atoms (metals, isotopes) encapsulated in the CNDs will permit 

the application of different imaging techniques (e.g., film, SPECT, PET). Because these CNDs can 

encapsulate a great number (2–1024) 6,39 of radionuclides per macromolecule, they can deliver at 

least a log fold more radioactivity to tumors than is possible with current radioimmunotherapeutic 

technologies. This offers the added potential of imaging microscopic tumor burdens as seen in 

micrometastases or minimal residual disease states. In the future, this amplified signal and molecular 

targeting capability could greatly aid in advancing the field of molecular imaging. 106 

Significantly, these composite nanodevices could be used for a combined imaging and therapy 

of the tumors. 6,39,106 

Many attempts have been made to target radionuclides to tumors 107–114 with the best known 

approaches being the use of radiolabeled monoclonal antibody therapy or the use of surface 

chelators bound to dendrimers. A detailed review and critique of these approaches is beyond the 

scope of this chapter, and several useful reviews exist. 115–118 In the past, radiolabeled antibodies for 

direct tumor cell targeting have received wide interest for the purpose of receptor imaging. In 

general, however, the penetration of these macromolecules into tumor tissue has been problematic 

with typically only 0.001%–0.01% of the injected dose being localized to each gram of solid human 

tumor tissue in man. Several factors account for this poor penetration: antibodies in the circulation 

must travel across the endothelial cell layer and often through dense fibrous stroma before encountering 

tumor cells; the dense packing of tumor cells and tight junctions between epithelial tumor 

cells hinders transport of the antibody within the tumor mass; the absence of lymphatics within the 

tumor contributes to the buildup of a high interstitial pressure that opposes the influx of molecules 

into the tumor core; antibodies entering the tumor become absorbed to the perivascular regions by 

the first tumor cells encountered, leaving none to reach tumor cells at sites farther from the blood 

vessels; and subpopulations of the rapidly mutating tumors can readily lose antigens targeted by the 

antibodies. Therefore, researchers have turned toward the development of small radiolabeled 

peptides that do not suffer from most of the aforementioned drawbacks. 107–118 

Tumor-directed antibody (or peptide) therapy has also not been able to deliver sufficient 

targeted dose because only one radioactive molecule can often be linked to a given antibody (or 

peptide), using a caged carrier to hold the radioactivity (e.g., DOTA). 119 The specific activity of the 

single radioactive moiety is also limited as it can damage the antibody if it is too high. This limits 

the detection of microscopic tumors, and higher doses leading to better detection have not been 

possible. This has also greatly limited any combination of imaging and therapy (where much higher 

dose delivery is needed). This derives at least partly from the fundamental problem that, to increase 

activity, one must increase substitution that tends to modify or obliterate the specificity of the 

antibody or targeting dendrimer. Finally, the ability to carry out different forms of imaging (PET or 

SPECT for example) is also limited by the chemistry necessary to place different isotopes onto the 

same targeting antibody. These fundamental problems can be overcome by the use of CNDs to 

deliver radioisotopes. 119 

Some of the fundamental problems that can be overcome by the use of targeted CNDs are those 

of solubility and dose delivery. First, many radioisotopes are often practically insoluble in water or 

body fluids, but they may become readily available in CNDs, i.e., in different sizes, contents, and 

surfaces. Second, use of CNDs permits at least a log-fold higher delivery of radioactivity than that 

available with current radioactive antibody therapies. As opposed to labeled antibodies that can 

carry only one radioisotope, 1–1000 nuclei can be encapsulated in one CND 6 , allowing the delivery 

of therapeutic dose amounts to tumor cells. 

To justify the targeted dose delivery by the any approach, it has to be first demonstrated that the 

treatment will be effective if the dose is delivered directly to the tumor (nanobrachytherapy). A 

proof-of-concept research of the therapeutic efficacy on { 198 Au} radioactive composite nanodevices 

in tumors (conceptually, other nuclides with the desired specific activity could also be used in 

the future, allowing the use of gamma, beta, alpha, or combinations of these radiations) has been 

conducted. To take advantage of the flexibility of composite nanodevice systems, Systemically 


584 

Targeted RadioTherapy (StaRT) is being developed to target radioactive nanodevices either to 

tumors, to the tumor angiogenic microvasculature, or to both for the treatment of cancer. Before 

carrying out these targeted experiments, however, an idea of the level of doses that could be 

delivered by CNDs and if these doses were enough to slow or halt the growth of tumors was 

needed. In order to test this idea of nanobrachytherapy, specific doses of radioactive CNDs were 

directly injected into growing tumors, and the dose needed to significantly slow tumor growth was 

examined. Once this dose is known, calculating if targeted devices intravascularly injected could 

deliver these types of doses can be conducted (i.e., if it is feasible to proceed with targeted 

nanodevice delivery). 

B16 melanoma carrying mice were randomly sorted into three groups of seven mice with 

approximately similar tumor volumes (w500 mm 3 ). The control group was intra-tumorally 

injected with PBS buffer, whereas the remaining two groups were intra-tumorally injected with 

35 and 74 mCi of radioactive 198 Au carrying composite nanodevice, respectively. A second control 

experiment was carried out with the same nanodevice after the radioactivity of its guest component 

decayed over time. The now cold {Au(0)} nanocomposite (corresponding to the original 74 mCi 

sample) was injected intra-tumorally into one set of tumor bearing mice, and the second was again 

injected with PBS buffer as control. The tumor volumes of all mice were recorded every other day 

for eight days. 

When compared to tumors injected with either cold nanocomposite or PBS (untreated), there 

was a statistically significant slowing of tumor growth observed in tumor injected with 74 mCi of 

radioactive gold nanocomposite (pZ0.03 in both cases). Although injection with the 35 mCi CND 

did not significantly slow tumor growth, there was a noticeable trend toward slow tumor growth 

(see Figure 28.24). It should be possible with the use of fractionated experiments (multiple injections 

over time) to be able to safely deliver these types of doses to tumors using targeted devices. 

With radioactive antibodies, intraperitoneal/intravenous injections of 300–800 mCi of various 

isotope/antibody combinations have been successfully utilized to slow tumor growth in mouse 

xenograft models. 120–123 It was also assumed that about 5%–10% of the 800 mCi injected dose 

(approximately 40–80 mCi) to the tumor could be targeted as supported by data on folate-surfaced 

targeted dendrimers 79 and the literature above on other targeted devices. With these calculations, 

the 35–74 mCi needed for tumor delivery from intravascular injections are feasible, especially if 

fractionated treatment is contemplated. 

Tumor Volume (mm 3 ) 

10000 

8000 

6000 

4000 

2000 

0 

(a) Radioactive Au Composite Nanodevice 

Treated Mice 

35 micro Ci 

74 micro Ci 

Untreated 

0 2 4 6 8 

Days 

Tumor Volume (mm 3 ) 

10000 

8000 

6000 

4000 

2000 

0 


(b) Cold Au Composite Nanodevice 

Treated Mice 

COLD 

Untreated 

0 2 4 6 8 

Days 

FIGURE 28.24 Treatment of melanoma tumors using radioactive Au-composite nanodevice. B16F10 melanoma 

tumors were grown in C57BL6/J mice to sizes of w500 mm 3 . (a) The mice were intra-tumorally injected 

with PBS buffer (untreated control) and 35 and 74 mCi of radioactive 198 Au carrying composite nanodevice 

(CND), respectively (nZ7 for all three groups). (b) Mice were intra-tumorally injected with PBS buffer 

(untreated control) and cold nanodevice corresponding to the original 74 mCi radioactive Au-CND (nZ3 

for both groups). The tumor volumes of all mice were recorded every other day for eight days. 



These experiments demonstrate for the first time that radiation therapy (nanobrachytherapy or 

systemic STaRT) with targeted composite nanodevices is feasible, and experiments are now being 

conducted to move this entire field forward with more detailed determination of the appropriate 

fractionation schemes for maximal tumor control using targeted composite nanodevices. Direct 

intratumoral injection gave a rough idea of the doses needed to do this, and they seemed reasonable 

given prior radioactive antibody data currently available in the literature. Importantly, for some 

tumors, direct injection of radioactive nanodevices may itself be a treatment. 


This research has been funded with federal funds in part from the U.S. Department of Energy under 

Award No. FG01-00NE22943 in part from the National Cancer Institute and National Institutes of 

Health, under Contract No. NO1-CO-97111 and RO1 CA104479 and in part by the Department of 

Defense Grant# DAMD17-03-1-0018. Silver nanocomposite research was sponsored by the Army 

Research Laboratory under the contract of DAAL-01-1996-02-0044. Thanks to D. Sorenson and C. 

Edwards of the Microscope and Imaging Laboratory at the University of Michigan for their 

technical support, to Leah Minc OSU-RC for INAA analysis, to Shraddha Nigavekar, Bindu 

Nair, Muhammed Taju Kariapper, Lok Yun Sung, Mikel Llanes, Kerstin May, Areej El-Jawari, 

Alla Kwitney, and Amber Warnat for their significant contributions on various components of the 

biologic components of these projects, and to Wojciech Lesniak, Kai Sun, X. Shi, and the 

University of Michigan Electron Microbeam Analysis Laboratory for their contributions to the 

synthesis and characterization components of these studies. 

DEFINITIONS 

Nanoscience The science of the phenomena related to nanoscale materials and/or devices. It refers 

to the multi-disciplinary study of materials, processes, and mechanisms that have distinctive 

properties (chemical, physical, and biological) from the molecular (i.e., !1 nm) to the 

sub-micrometer scales (i.e., O100 nm) because of quantum-size effects. 

Nanotechnology A wide range of possible technologies that encompass the incorporation of 

engineered or manufactured devices or materials with critical length scales between 

approximately 1 nm and 100 nm, and applications that exploit specific properties and 

functions of nanosized materials and/or devices. 

Particle Any object that has a persistent surface and an interior that may be different from 

the surface. 

Nanodevice A surface modified multifunctional nanoparticle that is able to perform 

different functions. 

Cluster Aggregations of atoms that are too large to be referred to as molecules and too small to 

resemble small pieces of crystals. Clusters generally do not have the same structure or 

atomic arrangement as a bulk solid and can change structure with the addition of just one 

or a few atoms. As the number of atoms (and their degree of freedom) increases, eventually 

a crystal-like structure may be established. Cluster science is devoted to understanding the 

changes in fundamental properties of materials, as a function of size, evolving from 

isolated atoms or small molecules to a bulk phase. 1,2 

Aggregates Dendrimers or nanoparticles held together by weak physical forces. 

Nomenclature To describe the complex structure of these nanoscopic complexes and nanocomposites, 

the following convention is used to describe the composition of the synthesized 

nanomaterials: brackets denote complexes, and braces represent nanocomposite 

structures. Within the brackets or braces, the complexed or encapsulated components 


586 

REFERENCES 


are listed (with an index of their average number per dendrimer molecule) followed by the 

family of dendrimers and the terms used for identification, i.e., core, generation, and 

surface. Naming of the materials will follow this pattern. For instance, 

PPI_DAB4.(NH 2) 16 describes a diaminobutane core poly(propyleneimine) dendrimer 

molecule with 16 primary amine terminal groups, and PAMAM_E5.(NH2)44(NHOAc)66 

denotes a generation five ethylenediamine core poly(amidoamine) dendrimer material 

with an average of 44 primary amine and 66 acetamide terminal groups (Please observe 

that the sum of these numbers are less than the theoretical 128, indicating that these are 

measured average numbers). All known properties can be listed by using this principle. 

Nanocomposites are also named according to the same pattern {(component#1)i(component#2)j.-FAMILY_Core.Generation.Terminal 

group}Therefore, the formula 

{(Au 0 ) 14.5-PAMAM_E4.NH 2} denotes a gold dendrimer nanocomposite where a generation 

4 amine terminated ethylenediamine (EDA or E for short) core poly(amidoamine) 

(PAMAM) dendrimer contains 14.5 zerovalent gold atoms per dendrimer molecule on 

average. The above scheme provides a relatively simple, but consistent, way to identify 

materials with a complicated structure. Additional descriptors may be added to the superscripts/superscripts 

and parentheses can be embedded into each other, e.g., 

fðAu 0 Þ10-E5:NH2g250G10 The composite listed above denotes an aggregate particle composed of 250G10 primary 

composite nanoparticles. When it is trivial, that dendrimer family is used; one may follow 

the shorter pattern and omit naming the family. Therefore, {(Au 0 )10-PAMAM_E5.NH2} 

and {(Au 0 ) 10-E5.NH2} may be considered to be equivalent names if only PAMAMs 

are used. 

1. Alivisatos, A. P. and Schultz, P. G., Organization of ‘nanocrystal molecules’ using DNA, Nature, 

382 (6592), 609–611, 1996. 

2. Schmid, G., Large clusters and colloids. Metals in the embryonic state, Chem. Rev., 92 (8), 

1709–1727, 1992. 

3. Jemal, A., Tiwari, R. C., Murray, T., Ghafoor, A., Samuels, A., Ward, E., Feuer, E. J., and Thun, 

M. J., Cancer statistics, CA Cancer J. Clin., 54 (1), 8–29, 2004. 

4. Khan, M. K., Nigavekar, S. S., Minc, L. D., Kariapper, M. S., Nair, B. M., Lesniak, W. G., and 

Balogh, L. P., In vivo biodistribution of dendrimers and dendrimer nanocomposites—implications 

for cancer imaging and therapy, Technol. Cancer Res. Treat., 4 (6), 603–613, 2005. 

5. Bielinska, A., Eichman, J. D., Lee, I., Baker, J. R., and Balogh, L., Imaging {Au0-PAMAM} gold– 

dendrimer nanocomposites in cells, J. Nanopart. Res., 4, 395–403, 2002. 

6. Grohn, F., Bauer, B. J., Akpalu, Y. A., Jackson, C. L., and Amis, E. J., Dendrimer templates for the 

formation of gold nanoclusters, Macromolecules, 33 (16), 6042–6050, 2000. 

7. McDevitt, M. R., Ma, D., Lai, L. T., Simon, J., Borchardt, P., Frank, R. K., Wu, K., Pellegrini, V., 

Curcio, M. J., and Miederer, M., et al., Tumor therapy with targeted atomic nanogenerators, Science, 

294 (5546), 1537–1540, 2001. 

8. Newkome, G. R., Moorefield, C. N., and Vögtle, F., Eds., Dendritic Macromolecules: Concepts, 

Syntheses, Perspectives, Wiley-VCH, New York, 1996. 

9. Frechet, J. M. J. and Tomalia, D. A., Eds., Dendrimers and Other Dendritic Polymers, Wiley, New 

York, p. 647, 2001. 

10. Mansfield, M. L. and Klushin, L. I., Monte Carlo studies of dendrimer macromolecules, Macromolecules, 

26 (16), 4262–4268, 1993. 

11. Karatasos, K., Adolf, D. B., and Davies, G. R., Statics and dynamics of model dendrimers as studied 

by molecular dynamics simulation, J. Chem. Phys., 115, 5310–5318, 2001. 



12. Murat, M. and Grest, G. S., Molecular dynamics study of dendrimer molecules in solvents of varying 

quality, Macromolecules, 29 (4), 1278–1285, 1996. 

13. Welch, P. and Muthukumar, M., Tuning the density profile of dendritic polyelectrolytes, Macromolecules, 

31 (17), 5892–5897, 1998. 

14. Tomalia, D. A., Baker, H., Dewald, J., Hall, M., Kallos, G., Martin, S., Roeck, J., Ryder, J., and 

Smith, P., Polym. J. (Tokyo), 17 (1), 117–132, 1985. 

15. Grayson, S. M. and Fréchet, J. M. J., Convergent dendrons and dendrimers: From synthesis to 

applications, Chem. Rev., 101 (12), 3819–3868, 2001. 

16. Duncan, R. and Izzo, L., Dendrimer biocompatibility and toxicity, Adv. Drug Deliv. Rev., 57 (15), 

2215–2237, 2005. 


(PAMAM) starburst dendrimers, J. Biomed. Mater. Res., 30 (1), 53–65, 1996. 

18. Tomalia, D. A., Dewald, J. R., Hall, M. J., Martin, S.J., and Smith, P. B., First SPSJ International 

Polymer Conference, Kyoto, Japan, p. 65, 1984. 

19. Tomalia, D. A., Naylor, A. M., and Goddard, W. A., Starburst dendrimers: Molecular-level control 

of size, shape, surface chemistry, topology, and flexibility from atoms to macroscopic matter, 

Angew. Chem. Int. Ed. Engl., 29 (2), 138–175, 1990. 

20. D’Emanuele, A. and Attwood, D., Dendrimer–drug interactions, Adv. Drug Deliv. Rev., 57 (15), 

2147–2162, 2005. 

21. Svenson, S. and Tomalia, D. A., Dendrimers in biomedical applications—reflections on the field, 

Adv. Drug. Deliv. Rev., 57 (15), 2106–2129, 2005. 

22. Tomalia, D. A. and Dvornic, P. R., In Polymeric Materials Encyclopedia, Salamone, J. C., Ed., Vol. 

3, CRC Press, Boca Raton, FL, pp. 1814–1830, 1996. 

23. Bielinska, A., Kukowska-Latallo, J. F., Johnson, J., Tomalia, D. A., and Baker, J. R. Jr., Regulation 

of in vitro gene expression using antisense oligonucleotides or antisense expression plasmids transfected 

using starburst PAMAM dendrimers, Nucleic Acids Res., 24 (11), 2176–2182, 1996. 

24. El-Sayed, M., Kiani, M. F., Naimark, M. D., Hikal, A. H., and Ghandehari, H., Extravasation of 

poly(amidoamine) (PAMAM) dendrimers across microvascular network endothelium, Pharm. Res., 

18 (1), 23–28, 2001. 

25. Malik, N., Evagorou, E. G., and Duncan, R., Dendrimer-platinate: A novel approach to cancer 

chemotherapy, Anticancer Drugs, 10 (8), 767–776, 1999. 

26. Raduchel, B., Schmitt-Willich, S., Ebert, J., Frezel, T., Misselwitz, B., and Weinmann, H. J., 

Synthesis and Characterization of Novel Dendrimer-Based Gadolinium Complexes as MRI Contrast 

Agents for the Vascular System, American Chemical Society, Washington, DC, p. 278, 1998. 

27. Wilbur, D. S., Pathare, P. M., Hamlin, D. K., Buhler, K. R., and Vessella, R. L., Biotin reagents for 

antibody pretargeting. 3. Synthesis, radioiodination, and evaluation of biotinylated starburst dendrimers, 

Bioconjug. Chem., 9 (6), 813–825, 1998. 

28. Shi, X., Banyai, I., Islam, M. T., Lesniak, W., Davis, D. Z., Baker, J. R. J., and Balogh, L. P., 

Generational, skeletal and substitutional diversities in generation one poly(amidoamine) dendrimers, 

Polymer, 46, 3022–3034, 2005. 

29. Uppuluri, S., Swanson, D. R., Brothers, H. M., Piehler, L. T., Li, J., Meier, D. J., Hagnauer, G. L., and 

Tomalia, D. A., Tecto(dendrimer) core-shell molecules: Macromolecular tectonics for the systemic 

synthesis of larger controlled structure molecules, Polym. Mater. Sci. Eng., 80, 55–56, 1999. 

30. Uppuluri, S., Swanson, D. R., Brothers, H. M. I., Piehler, L. T., Li, J., Meier, D. J., H., G. L., Tomalia, 

D.A, Core-Shell Tecto(Dendrimers): A New Class of Regio-Specifically Cross-Linked Polymers, 

Polymers for Advanced Technologies (PAT), Tokyo, Japan, August 31–September 5, pp. 94–95, 1999. 

31. Quintana, A., Raczka, E., Piehler, L., Lee, I., Myc, A., Majoros, I., Patri, A. K., Thomas, T., Mule, J., 

and Baker, J. R. Jr., Design and function of a dendrimer-based therapeutic nanodevice targeted to 

tumor cells through the folate receptor, Pharm. Res., 19 (9), 1310–1316, 2002. 

32. Balogh, L., Swanson, D.R., Spindler, R., Tomalia., D.A., Formation and characterization of 

dendrimer-based water soluble inorganic nanocomposites, In Proceedings of the American Chemical 

Society, Division of Polymeric Materials Science and Engineering, Vol. 77, Los Vegas, VV, pp. 118– 

119, 1997. 

33. Tan, N. B., Balogh, L., and Trevino, S., Structure of metallo-organic nanocomposites produced from 

dendrimer complexes, Proc. ACS Polym. Mater.: Sci. Eng., 77, 120, 1997. 


588 


34. Balogh, L. and Tomalia, D. A., Dendrimer-templated nanocomposites. I. Synthesis of zero-valent 

copper nanoclusters, J. Am. Chem. Soc., 120, 7355–7356, 1998. 

35. Zhao, M., Sun, L., and Crooks, R. M., Preparation of Cu nanoclusters within dendrimer templates, 

J. Am. Chem. Soc., 120 (19), 4877–4878, 1998. 

36. Esumi, K., Suzuki, A., Aihara, N., Usui, K., and Torigoe, K., Preparation of gold colloids with UV 

irradiation using dendrimers as stabilizer, Langmuir, 14 (12), 3157–3159, 1998. 

37. Larré, C., Donnadieu, B., Caminade, A. M., and Majoral, J. P., Regioselective gold complexation 

within the cascade structure of phosphorus-containing dendrimers, Chem. Eur. J., 4 (10), 2031–2036, 

1998. 

38. Bosman, A. W., Janssen, H. M., and Meijer, E. W., About dendrimers: Structure, physical properties, 

and applications, Chem. Rev., 99, 1665–1688, 1999. 

39. Diallo, M., Balogh, L., Shafagati, A., Johnson, J. H. J., Goddard, W. A. I., and Tomalia, D. A., 2- 

Poly(amidoamine) dendrimers: A new class of high capacity chelating agents for Cu (II) ions, 

Environ. Sci. Technol., 33 (5), 820–824, 1999. 

40. Diallo, M. S., Christie, S., Swaminathan, P., Balogh, L., Shi, X., Um, W., Papelis, C., Goddard, 

W. A. III, and Johnson, J. H. Jr., Dendritic chelating agents. 1. Cu(II) binding to ethylene diamine 

core poly(amidoamine) dendrimers in aqueous solutions, Langmuir, 20 (7), 2640–2651, 2004. 

41. Balogh, L., Valuzzi, R., Laverdure, K. S., Gido, S. P., Hagnauer, G. L., and Tomalia, D. A., 

Formation of silver and gold dendrimer nanocomposites, J. Nanopart. Res., 1 (3), 353–368, 1999. 

42. Balogh, L., Laverdure, K. S. Gido, S. P., Mott A. G., Miller, M. J., Ketchel, B. P., Tomalia, D. A. 

Dendrimer–metal nanocomposites, In Organic/Inorganic Hybrid Materials, pp. 69–75, 1999. 

43. Hong, S., Bielinska, A. U., Mecke, A., Keszler, B., Beals, J. L., Shi, X., Balogh, L., Orr, B. G., Baker, 

J. R. Jr., and Banaszak Holl, M. M., Interaction of poly(amidoamine) dendrimers with supported 

lipid bilayers and cells: Hole formation and the relation to transport, Bioconjug. Chem., 15 (4), 

774–782, 2004. 

44. Martin, C. R., Science, 266, 1961–1966, 1994. 

45. Shaw, C. F. III, Gold-based therapeutic agents, Chem. Rev., 99, 2589–2600, 1999. 

46. El-Sayed, I. H., Huang, X., and El-Sayed, M. A, Selective laser photo-thermal therapy of epithelial 

carcinoma using anti-EGFR antibody conjugated gold nanoparticles, Cancer Lett., 2005. 

47. Hirsch, L. R., Stafford, R. J., Bankson, J. A., Sershen, S. R., Rivera, B., Price, R. E., Hazle, N. J., and 

West, J. L., Nanoshell-mediated near-infrared thermal therapy of tumors under magnetic resonance 

guidance, Proc. Natl Acad. Sci. U.S.A., 100 (23), 13549–13554, 2003. 


cancer imaging and therapy, Nano Lett., 5 (4), 709–711, 2005. 

49. Loo, C., Lin, A., Hirsch, L., Lee, M., Barton, J., Halas, N., West, J., and Drezek, R., Nanoshell-enabled 

photonics-based imaging and therapy of cancer, Technol. Cancer Res. Treat., 3, 33–40, 2004. 

50. Holm, H. H. III, The history of interstitial brachytherapy of prostatic cancer, Semin. Surg. Oncol., 13, 

431–437, 1997. 

51. Root, S. W., Andrews, G. A., Knieseley, R. M., and Tyor, M. P., The distribution and radiation 

effects of intravenously administered colloidal Au 198 in man, Cancer, 7, 856–866, 1954. 

52. Rubin, P. and Levitt, S. H., The response of disseminated reticulum cell sarcoma to the intravenous 

injection of colloidal radioactive gold, J. Nucl. Med., 5, 581–594, 1964. 

53. Flocks, R. H., Kerr, H. D., Elkins, H. B., and Culp, A., Treatment of carcinoma of the prostate by 

interstitial radiation with radio-active gold: A preliminary report, J. Urol., 68, 510–522, 1952. 

54. Loening, S. A., Gold seed implantation in prostate brachytherapy, Semin. Surg. Oncol., 13 (6), 

419–424, 1997. 

55. Hainfeld, J. F., Foley, C. J., Srivastava, S. C., Mausner, L. F., Feng, N. I., Meinken, G. E., and 

Steplewski, Z., Radioactive gold cluster immunoconjugates: Potential agents for cancer therapy, Int. 

J. Radiat. Appl. Instrum. Part B, Nucl. Med. Biol. (Oxford), 17 (3), 287–294, 1990. 

56. Paciotti, G. F., Myer, L., Weinreich, D., Goia, D., Pavel, N., McLaughlin, R. E., and Tamarkin, L., 

Colloidal gold: A novel nanoparticle vector for tumor directed drug delivery, Drug Deliv., 11 (3), 

169–183, 2004. 

57. Garcia, M. E., Baker, L. A., and Crooks, R. M., Preparation and characterization of dendrimer–gold 

colloid nanocomposites, Anal. Chem., 71 (1), 256–258, 1999. 

58. Tomalia, D. A., Balogh, L. P., US Patent No. 6,664,315B2, 2003. 



59. Balogh, L., Tomalia, D. A., and Hagnauer, G. L., Revolution of nanoscale proportions, Chem. 

Innovation, 30 (3), 19–26, 2000. 

60. Ottaviani, M. F., Valluzzi, R., and Balogh, L., Internal structure of silver-poly(amidoamine) 

dendrimer complexes and nanocomposites, Macromolecules, 35 (13), 5105–5115, 2002. 

61. Balogh, L., Bielinska, A., Eichman, J. D., Valluzzi, R., Lee, I., Baker, J. R., Lawrence, T. S., and Khan, 

M. K., Dendrimer nanocomposites in medicine, Chim. Oggi (Chem. Today), 20 (5), 35–40, 2002. 

62. Cotton F. A., Willinson G., Advanced Inorganic Chemistry, 4 ed., Wiley, New York, pp. 1052–1055, 

1980. 

63. Lee, W. I., Bae, Y., and Bard, A. J., Strong blue photoluminescence and ECL from OH-terminated 

PAMAM dendrimers in the absence of gold nanoparticles, J. Am. Chem. Soc., 126, 8358–8359, 2004. 

64. Esumi, K., Suzuki, A., Yamahira, A., and Torigoe, K., Role of poly(amidoamine) dendrimers for 

preparing nanoparticles of gold, platinum, and silver, Langmuir, 16 (6), 2604–2608, 2000. 

65. Balogh, L., Ganser, T. R., and Shi, X. In Characterization of Dendrimer–Gold Nanocomposite 

Materials, Materials Research Society Symposium Proceedings, Warrendale, PA, 2005; Sanchez, 

C., Schubert, U., Laine, R. M., Chujo, Y., Materials Research Society, Warrendale, PA, EE13.33, 

pp. 1–6, 2005. 

66. Esumi, K., Dendrimers for nanoparticle synthesis and dispersion stabilization, Top. Curr. Chem., 

227, 31–52, 2003. 

67. Seo, Y.-S., Kim, K. S., Shin, K., White, H., Rafailovich, M., Sokolov, J., Lin, B., Kim, H. J., Zhang, 

C., and Balogh, L., Morphology of amphiphilic gold/dendrimer nanocomposite monolayers, Langmuir, 

18 (15), 5927–5932, 2002. 

68. Tomalia, D. A., Baker, H., Dewald, J., Hall, M., and Kallos, G., Dendritic macromolecules: 

Synthesis of starburst dendrimers, Macromolecules, 19, 2466–2468, 1986. 

69. Uppuluri, S., Swanson, D. R., Piehler, L. T., Li, J., Hagnauer, G. L., and Tomalia, D. A., Core-shell 

tecto(dendrimers): I. Synthesis and characterization of saturated shell models, Adv. Mater., 12 (11), 

796–800, 2000. 

70. Li, J., Swanson, D. R., Qin, D., Brothers, H. M., Piehler, L. T., Tomalia, D., and Meier, D. J., 

Characterizations of core-shell tecto-(dendrimer) molecules by tapping mode atomic force 

microscopy, Langmuir, 15 (21), 7347–7350, 1999. 

71. Tomalia, D. A., Piehler, L. T., Durst, H. D., and Swanson, D. R., Partial shell-filled core-shell 

tecto(dendrimers): A strategy to surface differentiated nano-clefts and cusps, Proc. Natl Acad. 

Sci. U.S.A., 99 (8), 5081–5087, 2002. 

72. Shi, X., Patri, A. K., Lesniak, W., Islam, M. T., Zhang, C., Baker, J. R., and Balogh, L. P., Analysis of 

poly(amidoamine)-succinamic acid dendrimers by slab-gel electrophoresis and capillary zone 

electrophoresis, Electrophoresis, 26, 2960–2967, 2005. 

73. Islam, M. T., Shi, X., Balogh, L. P., and Baker, J. R., HPLC separation of different generations of 

poly(amidoamine) dendrimers modified with various terminal groups, Anal. Chem., 77, 2063–2070, 

2005. 

74. Shi, X., Lesniak, W., Islam, M. T., Muñiz, M. C., Balogh, L. P., and Baker, J. R., Comprehensive 

characterization of surface-functionalized poly(amidoamine) dendrimers with acetamide, hydroxyl, 

and carboxyl groups, Colloids Surf., A: Physicochem. Eng. Aspects, 272, 139–150, 2006. 

75. Shi, X., Bányai, I., Islam, M. T., Lesniak, W., Davis, D. Z., Baker, J. R., and Balogh, L. P., 

Generational, skeletal and substitutional diversities in generation one poly(amidoamine) dendrimers, 

Polymer, 46 (9), 3022–3034, 2005. 

76. Lesniak, W., Bielinska, A. U., Sun, K., Janczak, K. W., Shi, X., Baker, J. R., and Balogh, L. P., 

Silver/dendrimer nanocomposites as biomarkers: Fabrication, characterization, in vitro toxicity, and 

intracellular detection, Nano Lett., 5 (11), 2123–2130, 2005. 

77. Nigavekar, S. S., Sung, L. Y., Llanes, M., El-Jawahri, A., Lawrence, T. S., Becker, C. W., Balogh, L., 

and Khan, M. K., 3H dendrimer nanoparticle organ/tumor distribution, Pharm. Res., 21 (3), 476–483, 

2004. 


Paulus, W., and Duncan, R., Dendrimers: Relationship between structure and biocompatibility 

in vitro, and preliminary studies on the biodistribution of 125I-labelled polyamidoamine dendrimers 

in vivo, J. Control. Release, 65 (1–2), 133–148, 2000. 


590 


79. Kukowska-Latallo, J. F., Candido, K. A., Cao, Z., Nigavekar, S. S., Majoros, I. J., Thomas, T. P., 

Balogh, L. P., Khan, M. K., and Baker, J. R. Jr., Nanoparticle targeting of anticancer drug improves 

therapeutic response in animal model of human epithelial cancer, Cancer Res., 65 (12), 5317–5324, 

2005. 

80. Tang, M. X., Redemann, C. T., and Szoka, F. C., In vitro gene delivery by degraded polyamidoamine 

dendrimers, Bioconjug. Chem., 7, 703–714, 1996. 

81. Rosi, N. L. and Mirkin, C. A., Nanostructures in biodiagnostics, Chem. Rev., 105, 1547–1562, 2005. 

82. Burda, C., Chen, X., Narayanan, R., and El-Sayed, M. A., Chemistry and properties of nanocrystals 

of different shapes, Chem. Rev., 105, 1025–1102, 2005. 

83. Bonkhoff, H. and Remberger, K., Widespread distribution of nuclear androgen receptors in the basal 

cell layer of the normal and hyperplastic human prostate, Virchows Arch. A Pathol. Anat. Histopathol., 

422 (1), 35–38, 1993. 


Nature, 380 (6572), 364–366, 1996. 

85. Rajotte, D. and Ruoslahti, E., Membrane dipeptidase is the receptor for a lung-targeting peptide 

identified by in vivo phage display, J. Biol. Chem., 274 (17), 11593–11598, 1999. 

86. Pasqualini, R., Koivunen, E., Kain, R., Lahdenranta, J., Sakamoto, M., Stryhn, A., Ashmun, R. A., 

Shapiro, L. H., Arap, W., and Ruoslahti, E., Aminopeptidase N is a receptor for tumor-homing 

peptides and a target for inhibiting angiogenesis, Cancer Res., 60 (3), 722–727, 2000. 


vasculature in a mouse model, Science, 279 (5349), 377–380, 1998. 


circulating ligands, Nat. Biotechnol., 15 (6), 542–546, 1997. 

89. Burg, M. A., Pasqualini, R., Arap, W., Ruoslahti, E., and Stallcup, W. B., NG2 proteoglycan-binding 

peptides target tumor neovasculature, Cancer Res., 59 (12), 2869–2874, 1999. 

90. Ellerby, H. M., Arap, W., Ellerby, L. M., Kain, R., Andrusiak, R., Rio, G. D., and Krajewski, S., 

Anti-cancer activity of targeted pro-apoptotic peptides, Nat. Med., 5 (9), 1032–1038, 1999. http:// 

www.java/Propub/medicine/nm0999_1032.fulltextjava/Propub/medicine/nm0999_1032.abstract. 

91. Dmitriev, I., Krasnykh, V., Miller, C. R., Wang, M., Kashentseva, E., Mikheeva, G., Belousova, N., 

and Curiel, D. T., An adenovirus vector with genetically modified fibers demonstrates expanded 

tropism via utilization of a coxsackievirus and adenovirus receptor-independent cell entry 

mechanism, J. Virol., 72 (12), 9706–9713, 1998. 

92. DeNardo, S. J., Burke, P. A., Leigh, B. R., O’Donnell, R. T., Miers, L. A., Kroger, L. A., and 

Goodman, S. L., Neovascular targeting with cyclic RGD peptide (cRGDf-ACHA) to enhance 

delivery of radioimmunotherapy, Cancer Biother. Radiopharm., 15 (1), 71–79, 2000. 

93. Janssen, M. L., Oyen, W. J., Dijkgraaf, I., Massuger, L. F., Frielink, C., Edwards, D. S., Rajopadhye, 

M., Boonstra, H., Corstens, F. H., and Boerman, O. C., Tumor targeting with radiolabeled alpha(v)beta(3) 

integrin binding peptides in a nude mouse model, Cancer Res., 62 (21), 6146–6151, 2002. 

94. Ogawa, M., Hatano, K., Oishi, S., Kawasumi, Y., Fujii, N., Kawaguchi, M., Doi, R., Imamura, M., 

Yamamoto, M., Ajito, K. et al., Direct electrophilic radiofluorination of a cyclic RGD peptide for 

in vivo alpha(v)beta(3) integrin related tumor imaging(1), Nucl. Med. Biol., 30 (1), 1–9, 2003. 

95. Haubner, R., Wester, H. J., Weber, W. A., Mang, C., Ziegler, S. I., Goodman, S. L., Senekowitsch- 

Schmidtke, R., Kessler, H., and Schwaiger, M., Noninvasive imaging of alpha(v)beta3 integrin 

expression using 18F-labeled RGD-containing glycopeptide and positron emission tomography, 

Cancer Res., 61 (5), 1781–1785, 2001. 

96. Sivolapenko, G. B., Skarlos, D., Pectasides, D., Stathopoulou, E., Milonakis, A., Sirmalis, G., 

Stuttle, A., Courtenay-Luck, N. S., Konstantinides, K., and Epenetos, A. A., Imaging of metastatic 

melanoma utilising a technetium-99m labelled RGD-containing synthetic peptide, Eur. J. Nucl. 

Med., 25 (10), 1383–1389, 1998. 

97. Liu, S., Edwards, D. S., Ziegler, M. C., Harris, A. R., Hemingway, S. J., and Barrett, J. A., 99mTclabeling 

of a hydrazinonicotinamide-conjugated vitronectin receptor antagonist useful for imaging 

tumors, Bioconjug. Chem., 12 (4), 624–629, 2001. 

98. Folkman, J., Tumor Angiogenesis, In Cancer Medicine, Holland, J. F., Bast, R. C., Morton, D. L., 

Frei, E., Kufe, D. W., and Weichselbaum, R. R., Eds., Vol. 1, Williams and Wilkens, Baltimore, MD, 

pp. 181–204, 1996. 



99. Jain, R. K., Determinants of tumor blood flow: A review, Cancer Res., 48 (10), 2641–2658, 1998. 

100. Ribatti, D., Vacca, A., and Dammacco, F., The role of the vascular phase in solid tumor growth: A 

historical review, Neoplasia, 1 (4), 293–302, 1999. 

101. Schlingemann, R. O., Rietveld, F. J., Kwaspen, F., van de Kerkhof, P. C., de Waal, R. M., and Ruiter, 

D. J., Differential expression of markers for endothelial cells, pericytes, and basal lamina in the 

microvasculature of tumors and granulation tissue, Am. J. Pathol., 138 (6), 1335–1347, 1991. 



smancs, Cancer Res., 46 (12 Pt 1), 6387–6392, 1986. 

103. Duncan, R., Malik, N., Richardson, S., and Ferruti, P., Polymer conjugates for anti-cancer agent and 

DNA delivery, Polym. Prepr. (Am. Chem. Soc. Div., Polym. Chem.), 39, 180, 1998. 

104. Muggia, F. M., Doxorubicin–polymer conjugates: Further demonstration of the concept of enhanced 

permeability and retention, Clin. Cancer Res., 5 (1), 7–8, 1999. 

105. Hainfeld, J. F., Foley, C. J., Srivastava, S. C., Mausner, L. F., Feng, N. I., Meinken, G. E., and 

Steplewski, Z., Radioactive gold cluster immunoconjugates: Potential agents for cancer therapy, Int. 

J. Rad. Appl. Instrum. B, 17 (3), 287–294, 1990. 

106. Sharma, V., Luker, G. D., and Piwnica-Worms, D., Molecular imaging of gene expression and 

protein function in vivo with PET and SPECT, J. Magn. Reson. Imaging, 16 (4), 336–351, 2002. 

107. Berning, D. E., Katti, K. V., Volkert, W. A., Higginbotham, C. J., and Ketring, A. R., 198Au-labeled 

hydroxymethyl phosphines as models for potential therapeutic pharmaceuticals, Nucl. Med. Biol.,25 

(6), 577–583, 1998. 

108. Brechbiel, M. W., Gansow, O. A., Wu, C., Garmestani, K., Yordanov, A., Deal, K., and Chappell, L., 

In Chelated Metal Ions for Therapeutic and Diagnostic Applications, 215th ACS Natioinal Meeting, 

Dallas, March 29–April 2, ACS, Washington, DC, Dallas, NUCL-106, 1998. 

109. Wiener, E. C., Brechbiel, M. W., Brothers, H., Magin, R. L., Gansow, O. A., Tomalia, D. A., and 

Lauterbur, P. C., Dendrimer-based metal chelates: A new class of magnetic resonance imaging 

contrast agents, Magn. Reson. Med., 31 (1), 1–8, 1994. 

110. Wu, C., Brechbiel, M. W., Kozak, R. W., and Gansow, O. A., Metal–chelate–dendrimer–antibody 

constructs for use in radioimmunotherapy and imaging, Bioorg. Med. Chem. Lett., 4 (3), 449–454, 

1994. 

111. Wiener, E. C., Konda, S., Shadron, A., Brechbiel, M., and Gansow, O., Targeting dendrimer-chelates 

to tumors and tumor cells expressing the high-affinity folate receptor, Investig. Radiol., 32 (12), 

748–754, 1997. 

112. Brechbeil, M., J. Nucl. Med., 35 (5 suppl. S), 62, 1994. 

113. Barth, R. F., Soloway, A. H., and Brugger, R. M., Boron neutron capture therapy of brain tumors: 

Past history, current status, and future potential, Cancer Investig., 14 (6), 534–550, 1996. 

114. Margerum, L. D., Campion, B. K., Koo, M., Shargill, N., Lai, J., Marumoto, A., and Sontum, P. C., 

J. Alloys Compd., 249 (1–2), 185–190, 1997. 

115. Von Kleist, S., Ten years of tumor imaging with labelled antibodies, In Vivo, 7 (6B), 581–584, 1993. 

116. Blankenberg, F. G. and Strauss, H. W., Nuclear medicine applications in molecular imaging, 

J. Magn. Reson. Imaging, 16 (4), 352–361, 2002. 

117. Kwekkeboom, D., Krenning, E. P., and de Jong, M., Peptide receptor imaging and therapy, J. Nucl. 

Med., 41 (10), 1704–1713, 2000. 

118. Ang, E. S. and Sundram, F. X., Prospects for tumour imaging with radiolabelled antibodies, Ann. 

Acad. Med. Singapore, 22 (5), 776–784, 1993. 

119. Deshpande, S. V., DeNardo, S. J., Kukis, D. L., Moi, M. K., McCall, M. J., DeNardo, G. L., and 

Meares, C. F., Yttrium-90-labeled monoclonal antibody for therapy: Labeling by a new macrocyclic 

bifunctional chelating agent, J. Nucl. Med., 31 (4), 473–479, 1990. 

120. Beaumier, P. L., Venkatesan, P., Vanderheyden, J. L., Burgua, W. D., Kunz, L. L., Fritzberg, A. R., 

Abrams, P. G., and Morgan, A. C. Jr., 186Re radioimmunotherapy of small cell lung carcinoma 

xenografts in nude mice, Cancer Res., 51 (2), 676–681, 1991. 

121. Chalandon, Y., Mach, J. P., Pelegrin, A., Folli, S., and Buchegger, F., Combined radioimmunotherapy 

and chemotherapy of human colon carcinoma grafted in nude mice, advantages and 

limitations, Anticancer Res., 12 (4), 1131–1139, 1992. 


592 


122. Cheung, N. K., Landmeier, B., Neely, J., Nelson, A. D., Abramowsky, C., Ellery, S., Adams, R. B., 

and Miraldi, F., Complete tumor ablation with iodine 131-radiolabeled disialoganglioside GD2specific 

monoclonal antibody against human neuroblastoma xenografted in nude mice, J. Natl 

Cancer Inst., 77 (3), 739–745, 1986. 

123. Esteban, J. M., Schlom, J., Mornex, F., and Colcher, D., Radioimmunotherapy of athymic mice 

bearing human colon carcinomas with monoclonal antibody B72.3: Histological and autoradiographic 

study of effects on tumors and normal organs, Eur. J. Cancer Clin. Oncol., 23 (6), 

643–655, 1987. 

124. Balogh,, L. P., Nigavekar, S. S., Cook, A. C., Minc, L., and Khan, M. K., Development of 

dendrimer–gold radioactive nanocomposites to treat cancer microvasculature, PharmaChem, 2 

(4), 94–99, 2003. 


29 

CONTENTS 

Applications of Liposomal 

Drug Delivery Systems 

to Cancer Therapy 


29.1 Introduction ....................................................................................................................... 595 

29.2 The Challenge of Cancer Therapy ................................................................................... 596 

29.3 The Rationale for the Use of Liposomal Drug Carriers .................................................. 597 

29.4 Differential Factors on the Pharmacokinetics of Low-Molecular- 

Weight Drugs and Liposomes .......................................................................................... 597 

29.5 Correlation Between Liposome Formulation and Liposome 

Pharmacokinetics-Stealth Liposomes ............................................................................... 598 

29.6 Preclinical Observations with Liposomal Anthracyclines ............................................... 600 

29.7 Liposomal Anthracyclines in the Clinic—Doxil (PLD) .................................................. 602 

29.8 Development of Other Liposome-Entrapped Cytotoxic Agents...................................... 605 

29.9 Future Avenues of Liposome Development—Receptor 

Targeted Liposomes.......................................................................................................... 606 

29.10 Conclusions ....................................................................................................................... 607 

References.................................................................................................................................... 607 


In the last two decades, major advances in the use of injectable drug delivery systems for the 

treatment of cancer have occurred. These include macromolecular conjugates, liposomes, and other 

nanoparticles. Poly(ethylene glycol) (PEG)-modified liposomal doxorubicin (Doxil, Caelyx) was 

the first liposomal anti-cancer drug to be approved by the Food and Drug Administration, 1 whereas 

paclitaxel albumin-bound particle suspension (ABI007, Abraxane) was recently approved for the 

treatment of metastatic breast cancer. 2 The ultimate goal behind the development of these sophisticated 

drug delivery systems is to improve the efficacy and decrease the side effects of new and old 

anti-cancer drugs. To achieve this goal, changes in the pharmacokinetics, biodistribution, and 

bioavailability profile leading to a positive impact on the drug pharmacodynamics are required. 

This chapter will focus on injectable liposome-based drug delivery systems of anti-cancer 

drugs, the most widely used drug nanoparticle in cancer. Liposomes are vesicles with an 

aqueous interior surrounded by one or more concentric bilayers of phospholipids with a diameter 

ranging from a minimal diameter of w30 nm to several microns. 3 However, for injectable 

clinical applications, practically all liposome formulations are in the submicron ultrafilterable 


595

596 

range (!200 nm size) and can be considered as nanosize particulate systems. Liposomes are 

spontaneously formed when amphiphilic lipids such as phospholipids are dispersed in water. 

The ensuing structures are physically stable supramolecular assemblies, and, unlike polymerized 

particles, they are not covalently bound. Whether the drug is encapsulated in the aqueous core or in 

the surrounding bilayer of the liposome is dependent on the characteristics of the drug and the 

encapsulation process. 4 In general, water-soluble drugs are encapsulated within the central aqueous 

core, whereas lipid-soluble drugs are incorporated directly into the lipid membrane. 

The current trend is to classify liposomes into a class of pharmaceutical devices in the nanoscale 

range engineered by physical and/or chemical means, and referred to as nanomedicines. 5 The 

field of nanomedicines is rapidly evolving and aims at increased sophistication in the design of 

nanosize devices and their interactions with cellular targets at the nanoscale level. In fact, liposomes 

are the first generation of nanomedicines approved for treatment of cancer (Doxil, described 

later in this chapter) and fungal infections (Ambisome, containing amphotericin B). Current liposome 

formulations involve a slow drug release system and a passive targeting process known as 

enhanced permeability and retention (EPR) that will be discussed later in this chapter. 

29.2 THE CHALLENGE OF CANCER THERAPY 


Current understanding of the molecular processes underlying the pathologic behavior of cancer 

cells has progressed enormously in the last decade. 6 Of particular relevance to cancer targeting is 

the fact that a number of receptors, mostly growth factor receptors, have been found to be overexpressed 

in tumor cells and to play an important role as catalysts of growth. Receptor profiling of 

tumors may offer a potential Achilles heel for targeting specific ligands or antibodies with or 

without delivery of a cytotoxic drug cargo. 7 In addition, the pathophysiology of tumor neovasculature 

and the interaction of tumor with stroma have been recognized as processes that play a major 

role in tumor development. Cancer is ultimately a disease caused by somatic gene mutations that 

result in the transformation of a normal cell into a malignant tumor cell. Eventually, the tumor cell 

phenotype progresses along three major steps: 8 

1. Increased proliferation rate and/or decreased apoptosis, causing an increase of tumor 

cell mass. 

2. Invasion of surrounding tissues and switch on of angiogenesis. This is a critical step that 

differentiates in situ, non-invasive tumors with no metastatic potential from invasive 

tumors with metastatic and life-threatening potential. 

3. Metastases, i.e., abnormal migration of tumor cells from the primary tumor site via blood 

vessels or lymphatics to distant organs with formation of secondary tumors. Most 

commonly, this is the process that causes death of the host because of disruption of 

the function of vital organs or systems (brain, lung, liver, kidney, bone marrow, coagulation, 

intestinal passage, and other). 

Despite formidable advances in clinical imaging, the diagnosis of a tumor mass requires usually 

the presence of a nodule of w10 mm diameter representing a cluster of 10 9 cells.* Because the 

lethal tumor burden is in the order of 10 12 cells in most cancer patients, this implies that tumors 

have already gone through 3/4 of their doubling cell expansion process by the time of clinical 

diagnosis. As a result, significant heterogeneity and phenotypic diversity are already present in 

* This does not apply to superficial skin tumors that can be recognized often when they are only 2–3 mm diameter and 

contain clusters of w10 7 cells. Occasionally, modern imaging techniques (high resolution CT scan, MRI) can detect deepseated 

lesions with suspected cancer features that are smaller than 10 mm. 


Applications of Liposomal Drug Delivery Systems to Cancer Therapy 597 

most diagnosed cancers, posing a major therapeutic challenge as a result of the development of 

metastatic ability and drug resistance. 

In addition to the lack of specificity of chemotherapeutic (cytotoxic) agents, a number of 

physiologic factors can seriously limit the efficiency of drug distribution from plasma to tumors 

and neutralize their effects. These include competition for drug uptake of well-perfused tissues such 

as liver and kidneys, rapid glomerular filtration and urinary excretion of low-molecular-weight 

drugs, protein binding with drug inactivation (e.g., cisplatin), and stability problems in biological 

fluids (e.g., hydrolysis of nitrosoureas, opening of lactone ring of camptothecin analogs). 

29.3 THE RATIONALE FOR THE USE OF LIPOSOMAL DRUG CARRIERS 

The rationale for the use of liposomes in cancer drug delivery is based on the following pharmacological 

principles: 1 

1. Slow drug release: drug bioavailability depends on drug release from liposomes. Entrapment 

of drug in liposomes will slow down drug release and reduce renal clearance to a 

variable extent. Slow release may range from a mere blunting of the peak plasma levels 

of free drug to a sustained release of drug, mimicking continuous infusion. These pharmacokinetic 

changes may have important pharmacodynamic consequences with regard 

to toxicity and efficacy of the liposome delivered agents. 

2. Site avoidance of specific tissues: the biodistribution pattern of liposomes may lead to a 

relative reduction of drug concentration in tissues specifically sensitive to the delivered 

drug. This may have implications with regard to the therapeutic window of various 

cytotoxic drugs such as the cardiotoxic anthracyclines, provided that anti-tumor efficacy 

is not negatively affected. 

3. Accumulation in tumors: prolongation of the circulation time of liposomes results in 

significant accumulation in tissues with increased vascular permeability. This is often the 

case of tumors, especially in those areas with active neoangiogenesis. 9 Tumor localization 

of long-circulating liposomes such as pegylated liposomes, sometimes referred to 

as Stealth or sterically-stabilized, 10 is a passive targeting effect, enabling substantial 

accumulation of liposome-encapsulated drug in the interstitial fluid at the tumor site, 11 

a phenomenon sometimes referred as EPR effect. 

29.4 DIFFERENTIAL FACTORS ON THE PHARMACOKINETICS OF LOW- 

MOLECULAR-WEIGHT DRUGS AND LIPOSOMES 

There are a number of differential effects of physiologic factors on clearance and biodistribution of 

low-molecular-weight drugs and nanoparticles such as liposomes. 

† Protein binding: low-molecular-weight drugs may be inactivated and/or irreversibly 

bound by plasma proteins, reducing the bioavailability toward cellular target molecules. 

This is the case of cisplatin, a widely used anti-cancer cytotoxic drug. In the case of 

nanoparticles, plasma proteins can adsorb to their surface, a process known as opsonization 

that results in tagging the particle for recognition and removal by macrophages. In 

addition, protein binding to the liposome surface may de-stabilize the bilayer and accelerate 

the leakage of liposome contents. PEG coating (PEGylation) of liposomes reduces 

opsonization and the effects associated with it. 

† Reticulo-endothelial system (RES) clearance: it is unimportant for low-molecularweight 

drugs, but it plays a major role in the clearance of nanoparticles, reducing the 


598 


TABLE 29.1 

Parameters Affecting Delivery of Liposomal Drugs to Tumors 

Tumor Factors Liposome Factors 

Blood flow Long circulation time 

Vascular permeability Stability (drug retention) 

Interstitial pressure Small vesicle size 

Phagocytic activity 

For a detailed discussion of these factors, see text. 

Saturation of the RES 

fraction available for distribution to tumor tissue. Kupffer cell macrophages lining the 

liver sinusoids remove opsonized liposomes and other nanoparticles from circulation and 

represent a major factor in the clearance of particulate carriers. 

† Glomerular filtration: unless they become protein-bound, low-molecular-weight drugs 

can be filtered out by kidney glomeruli. In contrast, liposomes and other nanoparticles 

are non-filterable because their radius exceeds the glomerular filterable threshold size. 

† Microvascular permeability: enhanced microvascular permeability with fenestrations in 

capillaries and post-capillary venules is critical for extravasation of nanoparticles from 

the blood stream to the interstitial fluid of target tissues. The presence of fenestrations is 

irrelevant for tissue delivery of small molecules. 

† Extravascular transport: diffusion is the predominant mechanism of transport for small 

molecules. In contrast, convective transport plays a major role in the extravascular 

movement of nanoparticles for which diffusion rates are very slow. 12 Large tumors 

tend to develop high interstitial pressure that reduces the rate of convective transport. 

In fact, in an animal model, it has been shown that liposomes accumulate significantly 

less in larger tumors on a per gram tissue basis. 13 In agreement with this, large tumor size 

predicts poor response to liposome-delivered chemotherapy in ovarian cancer. 14 

Table 29.1 lists a number of tumor and liposome factors that play an important role in the 

delivery of liposomal drugs. On the tumor side, a rich blood flow and highly permeable microvascular 

bed will increase the probability of liposome deposition, whereas a high interstitial fluid 

pressure is likely to reduce the movement of molecules and particles into the tumor compartment. 

On the liposome side, avoiding drug leakage and prolonging the circulation time will result in more 

liposomes reaching the tumor vascular bed with an intact drug payload, and a small vesicle size will 

facilitate extravasation through the endothelium gaps or fenestrations. There are also data indicating 

that saturation of the RES will prolong circulation time and indirectly enhance liposome 

deposition in tumors. 15 

29.5 CORRELATION BETWEEN LIPOSOME FORMULATION AND LIPOSOME 

PHARMACOKINETICS-STEALTH LIPOSOMES 

In 1971, Gregoriadis et al. 16 published the first research work where liposomes were used as drug 

carriers for medical applications. This initial study led to growing interest in liposomes, and many 

laboratories began examining liposome pharmacokinetics and biodistribution in animals as well as 

in vitro stability in serum. The early liposome work was mostly based on formulations composed of 

neutral egg lecithin (PC), often in combination with negatively or positively charged lipids. These 

liposomes were found to rapidly release a large fraction their encapsulated contents in circulation 

and were quickly removed from circulation by macrophages of the RES residing in liver and spleen. 



Reformulation with high phase-transition temperature (Tm) lipids (distearoyl-PC, dipalmitoyl-PC, 

sphingomyelin) and an addition of cholesterol led to improved retention of liposome contents 

and prolongation of circulation time, especially when the vesicles were properly downsized to 

!100 nm diameter. However, these relatively improved liposome formulations would still largely 

accumulate in the RES, and a greater improvement in circulation half-life appeared to be required 

for cancer targeting. Surface modifications of liposomes that could reduce the RES affinity were 

investigated based on the erythyrocyte paradigm whereby a layer of carbohydrate groups prolong 

circulation for nearly three months. A number of glycolipids such as monosialoganglioside (GM1), 

phosphatidyl-inositol, and cerebroside sulfate were included in the formulations and, indeed, 

extended liposome circulation time. 17,18 However, a major leap forward took place when the 

hydrophilic polymer PEG that was known to reduce immunogenicity and prolong circulation 

time when attached to enzymes and growth factors was introduced into liposomes in the early 

1990s. PEG that is inexpensive as a result of easy synthesis and that could be prepared in high purity 

and large quantities had distinct advantages over the other glycolipid surface modifiers. Addition of 

a conjugate of PEG with a lipid anchor, distearoyl-phosphatidylethanolamine (PEG–DSPE), to the 

liposomal formulation was shown to significantly prolong liposome circulation time 10,19–22 and 

formed a pivotal element of the pharmaceutical development of the DOXIL formulation described 

thereafter in this chapter. Because of their ability to avoid/escape RES clearance mechanisms, PEGcoated 

liposomes have been coined Stealth* liposomes. 23 In general, the pharmacokinetic variables 

of the liposomes depend on the liposomes’ physicochemical characteristics such as size, surface 

charge, membrane lipid packing, steric stabilization as conferred by PEG coating, dose, and route 

of administration. 

It should be noted that the pharmacokinetic disposition of liposome-associated agents is dependent 

on the carrier until the drug is released from the carrier. Therefore, the pharmacology and 

pharmacokinetics of drugs delivered in liposomes are complex, and ideally, studies must be done to 

evaluate the disposition of the entrapped form of the drug and of the released active drug. A 

prerequisite for these types of studies is to develop the proper methodology to discriminate 

between both forms of the drug in plasma. 

In parallel to the development of stable formulations with long circulation half-life, it was soon 

realized that a prolonged residence time in circulation was a critical pharmacokinetic factor for 

liposome deposition in tumors and that there was a strong correlation between liposome circulation 

time and tumor uptake. 18 A number of studies have addressed the mechanism of liposome accumulation 

in tumors. Microscopic observations with colloidal gold-labeled liposomes 24 and 

morphologic studies with fluorescent liposomes in the skin-fold chamber model 25 have demonstrated 

that liposomes extravasate into the tumor extracellular fluid through gaps in tumor 

microvessels and are found predominantly in the perivascular area with minimal uptake by 

tumor cells. Studies with ascitic tumors demonstrate a steady extravasation process of long circulating 

liposomes into the ascitic fluid with gradual release of drug followed by drug diffusion into 

the ascitic cellular compartment. 26,27 The process underlying the preferential tumor accumulation 

of liposomes as well as other macromolecular and particulate carriers is known as EPR effect. 28 

This is a passive and non-specific process, resulting from increased microvascular permeability and 

defective lymphatic drainage in tumors creating an in situ depot of liposomes in the tumor interstitial 

fluid. Circulating liposomes cross the leaky tumor vasculature, moving from plasma into the 

interstitial fluid of tumor tissue following convective transport and diffusion processes. Although 

convective transport of plasma fluid occurs also in normal tissues, the continuous, non-fenestrated 

endothelium and basement membrane prevents the extravasation of liposomes. EPR is relatively a 

slow process where long-circulating liposomes possess a distinct advantage because of the repeated 

* Stealth is a registered trademark of Alza Corp., Mountain View, CA. 


600 


passage through the tumor microvascular bed and their high concentration in plasma during an 

extended period of time. 

For any intra-vascular drug carrier device to access the tumor cell compartment and interact 

with tumor cell receptors, it must first cross the vascular endothelium and diffuse into the interstitial 

fluid because, with few exceptions, tumor cells and their surface receptors are not directly exposed 

to the blood stream. Therefore, the EPR effect is not only important for the tumor accumulation of 

non-targeted liposomes but also for that of ligand-targeted liposomes. This has led to the assertion 

that the extravasation process is the rate-limiting step of liposome accumulation in tumors. 29 

Experimental data with targeted and non-targeted liposomes have so far lent consistent support 

to this hypothesis. 30 

In most instances, delivery of drug to tumor cells depends on the release of drug from liposomes 

in the interstitial fluid because liposomes are seldom taken up by tumor cells unless they are tagged 

with specific ligands. The factors controlling this process and its kinetics are not well understood 

and may vary among tissues, depending on the liposome formulation in question. In the case of 

remote-loaded formulations, that is often the case of anthracyclines, a gradual loss of the liposome 

gradient retaining the drug in addition to disruption of the integrity of the liposome bilayer by 

phospholipases may be involved in the release process. Uptake by tumor-infiltrating macrophages 

could also contribute to liposomal drug release. In any case, once the drug is released from 

liposomes, it will freely diffuse through the interstitial tumor space and reach deep layers of 

tumor cells. This is an inherent advantage of this delivery system as opposed to covalently 

bound drug-carrier systems. It is also a critical factor for the success of the liposomal drug approach 

because most of the liposomes appear to remain in interstitial spaces immediately surrounding the 

blood vessels, 24,25 and therefore would not be able to interact with more than one layer of tumor 

cells. As a result, the ability of the liposome to effectively carry the anti-cancer agent to the tumor 

and the ability to release the drug at a satisfactory rate in the tumor extracellular fluid are equally 

important factors in determining the anti-tumor effect of liposomal-encapsulated anticancer 

agents. 

The EPR effect has been confirmed in a variety of implanted tumor models. Its validity 

regarding human tumors and, particularly, cancer metastases is unclear as yet. One point of 

concern is the fact that interstitial fluid pressure increases in most tumors once they grow above 

a certain size threshold, 31 therefore hindering extravascular transport and liposome delivery. Unfortunately, 

there is a paucity of imaging studies in cancer patients with radiolabeled liposomes. One 

of the few studies with radiolabeled pegylated liposomes demonstrated significant liposome 

accumulation based on tumor imaging findings in 15/17 patients tested. 32 In another study where 

tumor metastases and normal muscle tissues of two breast cancer patients were examined for 

doxorubicin concentration after injection of pegylated liposomal doxorubicin (PLD), liposomal 

drug was found at 10-fold greater concentration in tumor as compared to muscle. 33 Another 

important piece of work in this area is the study of Northfelt et al. 34 that pointed to an enhanced 

deposition of drug in Kaposi’s sarcoma (KS) skin lesions of patients receiving PLD as compared to 

normal skin of the same patients and to doxorubicin concentration in KS biopsies of patients 

receiving free doxorubicin. 

29.6 PRECLINICAL OBSERVATIONS WITH LIPOSOMAL ANTHRACYCLINES 

The drugs most frequently incorporated and evaluated in liposomal formulations are anthracyclines, 

including doxorubicin and daunorubicin. The choice of doxorubicin by many of the early 

research groups examining the role of liposomes as drug carriers in cancer chemotherapy stems 

from its broad spectrum of anti-tumor activity on the one hand, and its disturbing cumulative doselimiting 

cardiac toxicity. Anthracyclines such as doxorubicin and daunorubicin cause acute toxic 

side effects, including bone marrow depression, alopecia, and stomatitis, and they are dose limited 



by a serious and mostly irreversible characteristic cardiomyopathy. 35 The first study describing the 

encapsulation of anthracyclines into liposomes appeared in 1979. 36 Work from various research 

groups followed soon after, supporting the general principle that liposomal formulations reduced 

the toxicity of anthracyclines in animal models. 37 

Based on the Stealth concept and using an elegant drug-loading mechanism, a formulation of 

PLD known as DOXIL (Caelyx in Europe) has been developed (Figure 29.1). The loading 

mechanism, coined remote (active) loading, is based on an ammonium sulfate gradient, leading 

to marked accumulation of doxorubicin inside the aqueous phase of the liposome (w15,000 

doxorubicin molecules/vesicle). At high concentration, the entrapped drug forms a crystallinelike 

precipitate, contributing to the stability of the entrapment. 38,39 This loading technology provides 

substantial stability with negligible drug leakage in circulation while still enabling satisfactory rates 

of drug release in tissues and malignant effusions. 40 

Studies in animal tumor models with doxorubicin encapsulated in pegylated and other longcirculating 

liposomes have demonstrated increased anti-tumor activity of liposomal drug as 

compared to optimal doses of free drug in various rodent models of syngeneic and human 

tumors and increased accumulation of liposomal drug in various transplantable mouse and 

human tumors as compared to free drug with a delayed peak tumor concentration and slow 

tissue clearance after injection of liposomal drug. 40 

The most valuable pharmacokinetic advantage of the Stealth liposomal delivery system is the 

enhancement of tumor exposure to doxorubicin as a result of the accumulation of liposomes in 

tumors by the EPR effect (Figure 29.2) as demonstrated in animal models and in some forms of 

human cancer. When the tissue uptake of PLD was examined in a couple of syngeneic mouse tumor 

models, it was found that tumor drug uptake linearly correlated with dose, and liver drug uptake 

showed a saturation profile. In the case of free doxorubicin, liver uptake linearly increased with 

dose, and tumor uptake marginally increased with dose. As a result, the delta of tumor drug 

concentration in favor of PLD was substantially greater at high doses. 15 These results suggest 

FIGURE 29.1 Schematic diagram of pegylated (Stealth) liposome encapsulating doxorubicin in the water 

phase (DOXIL, PLD). The PEG coating is critical for stabilizing and protecting the liposome surface from 

opsonization by plasma proteins, conferring prolonged circulation time. The gradient-driven loading of doxorubicin 

results in a high drug concentration and gelification in the liposome water phase, ensuring stable drug 

retention during circulation. 


602 


FIGURE 29.2 (See color insert following page 522.) Deposition of PLD in M109 carcinoma subcutaneously 

implanted in BALB/c mice by the EPR effect. The mouse was sacrificed 48 h after intravenous injection of 

20 mg/kg PLD. Left panel: note the red-orange to bluish color in the tumor and peritumoral halo, suggesting 

accumulation of doxorubicin. Right panel: note the strong tumor fluorescence observed with UV light, 

confirming the presence of large amounts of doxorubicin. The drug levels measured in doxorubicinequivalents 

were: tumor, 86 mg/g (i.e., w35% of injected dose per gram); peritumoral halo, 40 mg/g; 

normal skin, 0.5 mg/g. 

a passive process of liposomal uptake into tumor with non-saturable kinetics, whereas the liver 

uptake is mediated by receptor-mediated endocytosis and shows a saturable profile. 

In preclinical therapeutic studies using a variety of rodent tumors and human xenografts in 

immunodeficient nude mice, PLD was more effective than free doxorubicin and other (non-pegylated) 

formulations of liposomal doxorubicin. 11 In a few instances, the activity of PLD preparations 

was matched, but not surpassed, by non-pegylated preparations of liposomal doxorubicin. 1 In many 

of these studies, the improved efficacy of PLD was obtained at milligram-equivalent doses to the 

MTD of free doxorubicin, indicating that there was a net therapeutic gain per milligram drug, 

independently of toxicity buffering. A study addressing this issue examined the activity of escalating 

doses of PLD and doxorubicin against implants of the mouse 3LL tumor and concluded that 

the activity of 1–2 mg/kg DOXIL was approximately equivalent to 9 mg/kg doxorubicin, i.e., a 

6-fold enhancement in efficacy. 41 Similar observations were made in the M109 model pointing to a 

4-fold advantage for PLD as compared to free doxorubicin, i.e., a dose of 2.5 mg/kg PLD was at 

least as effective as 10 mg/kg free doxorubicin. 15 

There is a large body of preclinical data on other liposome formulations of anti-cancer agents 

moving into clinical development or already approved for clinical use. A tentative list of these 

formulations is presented in Table 29.2. Obviously, in some cases, the added value of these 

formulations has not been or will not be sufficient to justify further development despite positive 

preclinical data. 

29.7 LIPOSOMAL ANTHRACYCLINES IN THE CLINIC—DOXIL (PLD) 

The anthracycline antibiotic doxorubicin has a broad spectrum of antineoplastic action and a 

correspondingly widespread degree of clinical use. In addition to its role in the treatment of 

breast cancer, doxorubicin is indicated in the treatment of various cancers of the lymphatic and 

hematopoyetic systems, gastric carcinoma, small-cell cancer of the lung, soft tissue, and bone 

sarcomas as well as cancer of the uterus, ovary, bladder, and thyroid. Unfortunately, toxicity 



TABLE 29.2 

Formulations of Liposomal Anti-Cancer Agents Clinically Tested a 

Pegylated Non-Pegylated Regional Therapy Non-Cytotoxic 

DOXIL (PLD) a 

PLD with DSPC 74,b 

MCC-465 

Immunoliposome 78 

SPI-077 63 

Lipoplatin 84 

Myocet 71 

DaunoXome 75 

Liposomal Annamycin 79 

DepoCyt (intrathecal) 72 

L-NDDP (intrapleural) 76 

Liposomal Camptothecin 

(aerosol) 80 

L-MTP-PE 73 

Liposomal ATRA 77 

BLP25 Liposomal Vaccine 81 

Liposomal Vincristine 57 

Liposomal IL2 (aerosol) 82 

Liposomal antisense ODN 83 

Lurtotecan/OSI-211 56,85 

Liposomal E1A (intratumoral) 

86 

S-CKD602 87,c 

OSI-7904L (TS inhibitor) 88 

LE-Paclitaxel 58 

a 

This list refers to formulations tested in the last 10 years and does not intend to be comprehensive given the difficulties 

involved in tracking all the relevant publications. 

b 

This formulation is from Taiwan and differs from DOXIL in that DSPC is the main lipid component. 

c 

Published information on this formulation of a topo I inhibitor is available only in abstract form. The reference given here 

refers to a preclinical study done with a similar formulation. 

often limits the therapeutic activity of doxorubicin and may preclude adequate dosing. Other 

common complications of conventional anthracycline therapy include alopecia and dose-limiting 

myelosuppression. Most importantly, cardiotoxicity limits the cumulative dose of conventional 

anthracycline that can be given safely. 42 

Encapsulation of anthracyclines within liposomes significantly alters the drug pharmacokinetic 

profile with a tendency to selective enhancement of drug concentration in tumors. 43 In animal 

studies, these pharmacologic effects resulted in maintained or enhanced anthracycline efficacy and 

safety in a variety of experimental tumor types. 44 Improved safety and/or therapeutic profile of 

liposomal anthracycline therapy in KS, ovarian cancer, breast cancer, and multiple myeloma have 

been reported. 45 The relative lack of cardiotoxicity with liposomal anthracycline therapy is an 

important asset of the liposomal approach. 42 Therefore, in most instances where conventional 

anthracycline therapy is effective but the required course of treatment could lead to a high risk 

of toxicity, the use of liposomal anthracyclines should be preferred. 

The most commonly used liposomal anthracycline formulation in the clinic is DOXIL 

(DOXIL w , Caelyx w , PLD). PLD embodies the pharmacokinetic advantages of stealth liposomes 

(Figure 29.1). The PLD formulation consists of doxorubicin encapsulated in pegylated liposomes 

formulated with a hydrogenated soybean PC and cholesterol. 40 PLD was granted initial market 

clearance in 1995 by the U.S. Food and Drug Administration (FDA) for use in treatment of AIDSrelated 

KS. In 1999, PLD was granted U.S. market clearance for use in the treatment of recurrent 

metastatic carcinoma of the ovary based on a superior therapeutic index over topotecan, the 

comparator drug. In January 2003, the European Commission of the European Union granted 

centralized marketing authorization to PLD as monotherapy for metastatic breast cancer based 

on equivalent efficacy to doxorubicin and reduced cardiac toxicity. In addition, a phase III trial 

recently completed in the U.S. and Europe investigating the safety and efficacy of PLD as a 

replacement for doxorubicin in multiple myeloma found equal efficacy and reduced toxicity for 

PLD treatment, paving the way for PLD to become standard therapy in multiple myeloma. 46 

PLD was recognized 10 years ago as a liposomal doxorubicin formulation with unique pharmacokinetics 

and had a major change in the clinical toxicity profile. Clinical pharmacokinetic 

studies have indicated that PLD dramatically prolongs the circulation time of doxorubicin 

in agreement with preclinical studies. In 1994, the results of a pharmacokinetic study where 


604 


15 patients were sequentially given the same dose in drug-equivalents of free doxorubicin and PLD 

were published. 47 A dramatic reduction in the drug clearance and volume of distribution resulting in 

a 1000-fold increase in the area under the curve (AUC) was observed with the liposomal formulation. 

It was also found that practically all the drug circulating in plasma is in liposomeencapsulated 

form. Metabolites in plasma were undetectable or were present at very low levels. 

However, they were readily detected in urine 24 hours or later after injection, indicating that the 

drug had become bioavailable. The following drug distribution picture had emerged from this 

initial study and more recent ones: 40 

1. Drug circulates in plasma for prolonged periods of time (half-life in the range of 

w50–80 hours) in liposome-encapsulated form. Despite its prolonged presence in 

blood, the drug is not bioavailable as long as it remains in the interior of a 

circulating liposome. 

2. Most of the injected drug (O95%) is distributed to tissues in liposome-encapsulated 

form. Once in tissues, drug leakage and liposome breakdown with or without liposome 

internalization by cells gradually provides a pool of bioavailable drug. 

3. The rate of metabolite production is slower than the rate of renal clearance of metabolites. 

As a result, metabolites do not accumulate in plasma but can be detected in urine. 

4. A small fraction of injected drug (!5%) leaks from circulating liposomes and is handled 

as free drug with fast plasma clearance and rapid metabolism. This drug fraction is the 

source of small amounts of metabolites that can be sometimes detected in plasma. 

In 1995, a phase I study of PLD in patients with solid tumors 48 provided clear evidence of a 

major change in the toxicity profile with muco-cutaneous toxicities as the major dose-limiting 

toxicities. In contrast, myelosuppression and alopecia were minor and cardiotoxicity was conspicuously 

absent. The maximal tolerated dose was established as 60 mg/m 2 with mucositis being the 

dose-limiting toxicity. It was also found that the optimal dosing interval for retreatment was four 

weeks, rather than the standard 3-week schedule of doxorubicin. The dose-schedule limiting 

toxicity was a form of skin toxicity known as hand-foot syndrome, also referred to as palmarplantar 

erythrodysesthesia (PPE), that appears to be related to the long half-life of PLD. Therefore, 

it became well-established that the PLD formulation imparts a significant pharmacokinetic–pharmacodynamic 

change to the drug doxorubicin, unprecedented in magnitude for any intravenous 

drug delivery system. Later, data gathered from further clinical studies in metastatic breast cancer 

and recurrent ovarian cancer brought down the recommended dose of PLD from 60 to 40–50 mg/m 2 

once every four weeks. 49 This dose reduction was needed mainly to prevent skin toxicity, resulting 

from successive courses of therapy. 

KS is a multi-focal tumor affecting skin and sometimes mucosas well known for its very high 

microvascular permeability. Profuse extravasation of colloidal gold-labeled Stealth liposomes in a 

transgenic mouse model of KS has been shown. 50 In AIDS patients, KS is frequent and has an 

aggressive course. Therefore, this condition was chosen for initial clinical testing of PLD in phase 

II-III studies. Indeed, PLD as single agent therapy demonstrated a significantly greater efficacy and 

better safety than standard chemotherapy (combinations of bleomycin and vincristine with or 

without doxorubicin) and was also effective as second line chemotherapy in pretreated patients. 

Because of the extreme sensitivity of KS to chemotherapy, a low and relatively subtoxic dose of 

20 mg/m 2 every three weeks is sufficient for effective treatment. 51 

Further to the successful application in various malignant conditions, there was another major 

development in the clinical research with PLD related to the prevention of the cardiac toxicity of 

doxorubicin. Evaluation of the cardiac function in patients receiving PLD revealed a major risk 

reduction of cardiotoxicity as compared to free doxorubicin historical data. A retrospective analysis 

of patients treated with large cumulative doses of PLD did not reveal any significant cardiac toxicity 



despite the fact that some of these patients were treated with three times as much as the maximal 

cumulative dose acceptable for free doxorubicin. 52 Two additional reports focusing on cardiac 

biopsies of PLD-treated patients at high cumulative doses demonstrated minimal pathology findings 

in line with the lack of clinical cardiotoxicity. 53,54 The cardiac safety of PLD was confirmed in 

a phase III study in metastatic breast cancer, showing a drastic reduction in cardiotoxicity in PLDtreated 

patients in comparison to free doxorubicin treatment. 55 

29.8 DEVELOPMENT OF OTHER LIPOSOME-ENTRAPPED 

CYTOTOXIC AGENTS 

Besides anthracyclines, other cytotoxic compounds currently being developed in liposomes 

include: Topoisomerase I inhibitory properties such as camptothecin analogs, 56 mitotic spindle 

poisons such as Vinca alkaloids and taxanes, 57,58 and DNA-reactive agents such as cisplatin, 59 and 

mitomycin C (MMC). 60 For cell cycle phase-specific cytotoxic drugs of the first two groups (e.g., 

topotecan, vinorelbine), the anti-tumor activity tends to increase with liposome encapsulation as the 

time of exposure is extended by exploiting the liposomal slow release features. An additional 

advantage for liposome encapsulation of camptothecin analogs derives from their improved stability 

in the mildly acidic environment of the liposome water phase as opposed to the instability of 

their active lactone configuration when present in free form at pH 7.4 in plasma. For cisplatin, a 

major advantage of liposome encapsulation would be reduction of the nephrotoxicity. At this time, 

none of these compounds have yet been approved by a regulatory agency and none has gone 

through phase III studies. 

It should be noted that the development of liposomal formulations of cytotoxic agents has often 

failed for different reasons. The formulation of a liposomal anti-cancer agent is a complex process 

with at least three distinct variables that may affect the outcome and the risk of failure: the design of 

the liposome carrier, the choice of the drug, and the method of drug encapsulation. One example 

illustrating the complexity of the relationship between formulation and pharmacological effects is 

provided by a product known as SPI-77, consisting of cisplatin encapsulated in pegylated liposomes. 

These liposomes have long-circulating characteristics, retain the drug in plasma 

exceedingly well, reduce drug toxicity, and produce a high and long-lasting accumulation of 

drug in implanted tumors. 61 However, because the Pt exposure was measured in tumor extracts, 

it is unclear whether the Pt measured was SPI-77 (i.e., liposomally encapsulated Pt), protein-bound 

Pt, or unbound Pt. The anti-tumor activity of SPI-77 in preclinical models was variable and, in some 

cases, inferior by comparison to free drug. 61 In clinical studies, SPI-77 was inactive and showed no 

dose-limiting toxicities even at doses greater than 2-fold the MTD of free cisplatin. 62 It turned out 

that cisplatin release from these liposomes was minimal in in vitro 61 as well as in vivo systems as 

indicated by the low occurrence of DNA adducts, resulting in a major reduction of bioavailability. 63 

Indeed, microdialysis experiments in animal tumor models suggest that the amount of unbound Pt 

released by SPI-77 in tumor extracellular fluid is significantly lower when compared with free 

cisplatin. 64 The lack of activity of SPI-77 in early-phase clinical studies has led to the discontinuation 

of its further clinical development. 

In some instances, a strategy based on chemically modifying the drug is needed to achieve 

stable encapsulation. It is a formidable challenge to develop liposome formulations that retain the 

drug payload in the blood stream during prolonged periods of circulation, thereby taking full 

advantage of the pharmacokinetic and biodistribution benefits of the pegylated liposomal carriers. 40 

Passive encapsulation of MMC in liposomes and even encapsulation of cyclodextrin–MMC 

complexes in liposomes did not result in stable formulations as a result of the rapid leakage of 

MMC. Recently, an alternative approach to delivering MMC via long-circulating liposomes has 

been proposed based on the design of a lipophilic prodrug with excellent retention in liposomes, but 

gradually releasing active MMC upon exposure to appropriate environmental conditions. 60 


606 


Attachment of drugs to bilayer-compatible lipids can ensure prolonged association with a liposome. 

65,66 The MMC prodrug consists of a conjugate of MMC attached to 1,2-distearoyl glycerol 

lipid via a cleavable dithiobenzyl linker. Upon thiolytic cleavage of the disulfide-substituted benzyl 

urethane, MMC is released and becomes bioavailable. Intravenous administration of a pegylated 

liposomal formulation of the MMC prodrug in rats resulted in long circulation time and slow 

clearance consistent with Stealth pharmacokinetics, indicating stable prodrug retention in liposomes. 

The therapeutic index of this liposomal formulation was superior to that of free MMC in 

three animal tumor models, including a mouse tumor model with a multidrug-resistant phenotype 

where PLD was essentially inactive. The potential added value of this approach is 2-fold: effective 

treatment of multidrug- resistant tumors while significantly buffering the toxicity of MMC. 

29.9 FUTURE AVENUES OF LIPOSOME DEVELOPMENT—RECEPTOR 

TARGETED LIPOSOMES 

As liposomes remain one of the most attractive platforms for systemic drug delivery, an increased 

expansion and sophistication of these systems would be expected in the forthcoming future. The 

most logical improvement is the coupling of a ligand to the surface of the liposome that will target 

the vesicles to a specific cell-surface receptor followed by internalization and intracellular delivery 

of the liposome drug cargo. Examples in this direction are the targeting of PLD to HER2-expressing 

or folate-receptor expressing cancer cells using, respectively, a specific anti-HER2 scFv 67 or a 

folate conjugate anchored to the liposome surface. 30 Another example is the tumor vascular 

targeting of liposomes with endothelium-specific peptides associated to liposomes. 68 Amajor 

advantage of targeted liposomal nanocarriers over ligand-drug bioconjugates is the delivery-amplifying 

effect of the former that sometimes can provide to the target cell a ratio of 1,000 drug 

molecules per single ligand–receptor interaction. 

One important technological and conceptual advance in this field is the ability to insert ligands 

onto preformed liposomes using lipophilic tails as anchors into the bilayer. Ligand post-insertion 

(as opposed to pre-insertion during liposome formation) has been demonstrated for antibody fragments 

69 and small ligands such as folate. 30 This approach enables, on the one hand, the possibility 

of using end-product formulations without interfering with the pharmaceutical methods of preparation, 

and, on the other hand, the flexibility of ligand choice, depending on the target to be 

addressed. One encouraging observation, at least for folate-targeted liposomes, is that pegylated 

liposomes are able to retain during in vivo circulation and extravasation to malignant ascites the 

targeting ligand and the ability to bind to target cells, suggesting that these systems are stable 

enough for in vivo applications, requiring prolonged circulation before encountering the target. 70 

Except for rare instances, tumor cells are not directly exposed to the blood stream. Therefore, 

for an intra-vascular targeting device to access the targeted tumor cell receptor, it has first to cross 

the vascular endothelium and diffuse into the interstitial fluid. Experimental data with antibodytargeted 

liposomes and folate-targeted liposomes indicate that liposome deposition in tumors is 

similar for both targeted and non-targeted systems, 29,70 supporting the hypothesis that extravasation 

is the rate-limiting step of liposome accumulation in tumors. However, once liposomes have 

penetrated the tumor interstitial fluid, binding of targeted liposomes to the targeted receptor may 

occur, shifting the intra-tumor distribution from the extracellular compartment to the tumor cell 

compartment as recently shown for a mouse ascitic tumor. 70 Retrograde movement of liposomes 

into the blood stream, if any, will be reduced for liposomes with binding affinity to a tumor cell 

receptor. Therefore, the theoretical advantages of a targeted liposomal drug delivery system over a 

non-targeted one are primarily derived from a relative shift of liposome distribution from the 

extracellular fluid to the tumor cell compartment, and, possibly, from prolonged liposome retention 

in tumor. On the negative side, the main disadvantage of a targeted system to a cancer cell receptor 

is the difficulty of a large nanosize assembly such as liposomes to penetrate a solid tumor mass, 



confronted by the high interstitial fluid pressure that often characterizes tumor masses at clinically 

detectable size. 31 One such system targeted to the Her2 receptor has shown superior therapeutic 

efficacy over its non-targeted counterpart. With the maturation of the pharmaceutical technology 

required for the application of targeted liposomal drug delivery systems, the final word regarding 

their added value over non-targeted systems awaits clinical trials. 


Despite huge research efforts, with liposomal drugs is cancer therapy anthracycline formulations, 

and specifically, the leading compound, PLD, are the only liposome products that have been able, 

so far, to fill a niche in the anti-cancer armamentarium. As this chapter has discussed, this is due, at 

least in part, to the complexity of the pharmaceutical technology involved that requires drugspecific 

solutions and has a direct impact on the pharmacology of the end compound. Many of 

the oncology opinion leaders regard liposome formulations, with liposomal in cancer therapy drugs 

as a cosmetic improvement of toxic drugs that will become obsolete in the near future once smart 

new drugs are developed, given the increasing knowledge of the molecular biology of cancer cells. 

However, a realistic scenario favored by the majority of clinicians predicts that the use of broadspectrum 

cytotoxic drugs will remain as a major tool in cancer therapy for a few more decades to 

come. Moreover, the nanomedicine approach is perceived today as a key to future breakthroughs 

because of the complexity of living systems. These factors may reinvigorate the interest in particulate 

drug delivery systems such as liposomes and augur that the use of liposome carriers in cancer 

therapy will expand in the foreseeable future. 

REFERENCES 


chemotherapy, Cancer Invest., 19(4), 424–436, 2001. 

2. Gradishar, W. J., Tjulandin, S., Davidson, N., Shaw, H., Desai, N., Bhar, P. et al., Phase III trial of 

nanoparticle albumin-bound paclitaxel compared with polyethylated castor oil-based paclitaxel in 

women with breast cancer, J. Clin. Oncol., 23(31), 7794–7803, 2005. 

3. Lasic, D. D. and Papahadjopoulos, D., Medical Applications of Liposomes, Elsevier, Amsterdam, 

1998. 

4. Barenholz, Y., Relevancy of drug loading to liposomal formulation therapeutic efficacy, J. Liposome 

Res., 13(1), 1–8, 2003. 

5. Moghimi, S. M., Hunter, A. C., and Murray, J. C., Nanomedicine: Current status and future prospects, 

FASEB J., 19(3), 311–330, 2005. 

6. Hanahan, D. and Weinberg, R. A., The hallmarks of cancer, Cell, 100(1), 57–70, 2000. 

7. Van Den Bossche, B. and Van de Wiele, C., Receptor imaging in oncology by means of nuclear 

medicine: Current status, J. Clin. Oncol., 22(17), 3593–3607, 2004. 

8. Kastan, M. B., Molecular biology of cancer: The cell cycle, In Cancer: Principles & Practice of 

Oncology, DeVita, V. T., Hellman, S., and Rosenberg, S. A., Eds., 5th ed., Lippincott-Raven, Philadelphia, 

pp. 121–134, 1997. 

9. Jain, R. K., Delivery of molecular medicine to solid tumors: Lessons from in vivo imaging of gene 

expression and function, J. Control. Release, 74(1-3), 7–25, 2001. 

10. Papahadjopoulos, D., Allen, T. M., Gabizon, A., Mayhew, E., Matthay, K., Huang, S. K. et al., 

Sterically stabilized liposomes: Improvements in pharmacokinetics and antitumor therapeutic efficacy, 

Proc. Natl Acad. Sci. USA, 88(24), 11460–11464, 1991. 

11. Gabizon, A. and Martin, F., Polyethylene glycol-coated (pegylated) liposomal doxorubicin. Rationale 

for use in solid tumours, Drugs, 54(Suppl. 4), 15–21, 1997. 

12. Swabb, E. A., Wei, J., and Gullino, P. M., Diffusion and convection in normal and neoplastic tissues, 

Cancer Res., 34(10), 2814–2822, 1974. 


608 


13. Harrington, K. J., Rowlinson-Busza, G., Syrigos, K. N., Abra, R. M., Uster, P. S., Peters, A. M. et al., 

Influence of tumour size on uptake of 111 ln-DTPA-labelled pegylated liposomes in a human tumour 

xenograft model, Br. J. Cancer, 83(5), 684–688, 2000. 

14. Safra, T., Groshen, S., Jeffers, S., Tsao-Wei, D. D., Zhou, L., Muderspach, L. et al., Treatment of 

patients with ovarian carcinoma with pegylated liposomal doxorubicin: Analysis of toxicities and 

predictors of outcome, Cancer, 91(1), 90–100, 2001. 

15. Gabizon, A., Tzemach, D., Mak, L., Bronstein, M., and Horowitz, A. T., Dose dependency of 

pharmacokinetics and therapeutic efficacy of pegylated liposomal doxorubicin (DOXIL) in murine 

models, J. Drug Target, 10(7), 539–548, 2002. 

16. Gregoriadis, G., Leathwood, P. D., and Ryman, B. E., Enzyme entrapment in liposomes, FEBS Lett., 

14(2), 95–99, 1971. 

17. Allen, T. M. and Chonn, A., Large unilamellar liposomes with low uptake into the reticuloendothelial 

system, FEBS Lett., 223(1), 42–46, 1987. 


blood and enhanced uptake by tumors, Proc. Natl Acad. Sci. USA, 85(18), 6949–6953, 1988. 

19. Allen, T. M., Austin, G. A., Chonn, A., Lin, L., and Lee, K. C., Uptake of liposomes by cultured 

mouse bone marrow macrophages: Influence of liposome composition and size, Biochim. Biophys. 

Acta, 1061(1), 56–64, 1991. 

20. Woodle, M. C. and Lasic, D. D., Sterically stabilized liposomes, Biochim. Biophys. Acta, 1113(2), 

171–199, 1992. 

21. Klibanov, A. L., Maruyama, K., Torchilin, V. P., and Huang, L., Amphipathic polyethyleneglycols 

effectively prolong the circulation time of liposomes, FEBS Lett., 268(1), 235–237, 1990. 

22. Allen, T. M., Hansen, C., Martin, F., Redemann, C., and Yau-Young, A., Liposomes containing 

synthetic lipid derivatives of poly(ethylene glycol) show prolonged circulation half-lives in vivo, 

Biochim. Biophys. Acta, 1066(1), 29–36, 1991. 

23. Lasic, D. D. and Martin, F. J., Stealth Liposomes, CRC Press, Boca Raton, FL, 1995. 

24. Huang, S. K., Lee, K. D., Hong, K., Friend, D. S., and Papahadjopoulos, D., Microscopic localization 

of sterically stabilized liposomes in colon carcinoma-bearing mice, Cancer Res., 52(19), 5135–5143, 

1992. 



xenograft, Cancer Res., 54(13), 3352–3356, 1994. 

26. Gabizon, A. A., Selective tumor localization and improved therapeutic index of anthracyclines 

encapsulated in long-circulating liposomes, Cancer Res., 52(4), 891–896, 1992. 

27. Bally, M. B., Masin, D., Nayar, R., Cullis, P. R., and Mayer, L. D., Transfer of liposomal drug carriers 

from the blood to the peritoneal cavity of normal and ascitic tumor-bearing mice, Cancer Chemother. 

Pharmacol., 34(2), 137–146, 1994. 

28. Maeda, H., The enhanced permeability and retention (EPR) effect in tumor vasculature: The key role 

of tumor-selective macromolecular drug targeting, Adv. Enzyme Regul., 41, 189–207, 2001. 

29. Goren, D., Horowitz, A. T., Zalipsky, S., Woodle, M. C., Yarden, Y., and Gabizon, A., Targeting of 

stealth liposomes to erbB-2 (Her/2) receptor: In vitro and in vivo studies, Br. J. Cancer, 74(11), 

1749–1756, 1996. 

30. Gabizon, A., Shmeeda, H., Horowitz, A. T., and Zalipsky, S., Tumor cell targeting of liposomeentrapped 

drugs with phospholipid-anchored folic acid-PEG conjugates, Adv. Drug Deliv. Rev., 56(8), 

1177–1192, 2004. 

31. Stohrer, M., Boucher, Y., Stangassinger, M., and Jain, R. K., Oncotic pressure in solid tumors is 

elevated, Cancer Res., 60(15), 4251–4255, 2000. 

32. Harrington, K. J., Mohammadtaghi, S., Uster, P. S., Glass, D., Peters, A. M., Vile, R. G. et al., 

Effective targeting of solid tumors in patients with locally advanced cancers by radiolabeled pegylated 

liposomes, Clin. Cancer Res., 7(2), 243–254, 2001. 

33. Symon, Z., Peyser, A., Tzemach, D., Lyass, O., Sucher, E., Shezen, E. et al., Selective delivery of 

doxorubicin to patients with breast carcinoma metastases by stealth liposomes, Cancer, 86(1), 72–78, 

1999. 



34. Northfelt, D. W., Martin, F. J., Working, P., Volberding, P. A., Russell, J., Newman, M. et al., 

Doxorubicin encapsulated in liposomes containing surface-bound polyethylene glycol: Pharmacokinetics, 

tumor localization, and safety in patients with AIDS-related Kaposi’s sarcoma, J. Clin. 

Pharmacol., 36(1), 55–63, 1996. 

35. Young, R. C., Ozols, R. F., and Myers, C. E., The anthracycline antineoplastic drugs, N. Engl. J. Med., 

305(3), 139–153, 1981. 

36. Forssen, E. A. and Tokes, Z. A., In vitro and in vivo studies with adriamycin liposomes, Biochem. 

Biophys. Res. Commun., 91(4), 1295–1301, 1979. 

37. Gabizon, A. A., Liposomal anthracyclines, Hematol. Oncol. Clin. North Am., 8(2), 431–450, 1994. 

38. Lasic, D. D., Frederik, P. M., Stuart, M. C., Barenholz, Y., and McIntosh, T. J., Gelation of liposome 

interior. A novel method for drug encapsulation, FEBS Lett., 312(2–3), 255–258, 1992. 

39. Haran, G., Cohen, R., Bar, L. K., and Barenholz, Y., Transmembrane ammonium sulfate gradients in 

liposomes produce efficient and stable entrapment of amphipathic weak bases, Biochim. Biophys. 

Acta, 1151(2), 201–215, 1993. 

40. Gabizon, A., Shmeeda, H., and Barenholz, Y., Pharmacokinetics of pegylated liposomal Doxorubicin: 

Review of animal and human studies, Clin. Pharmacokinet., 42(5), 419–436, 2003. 

41. Colbern, G. T., Hiller, A. J., Musterer, R. S. et al., Significant increase in antitumor potency of 

doxorubicin HCI by its encapsulation in pegylated liposomes, J. Liposome Res., 9, 523–538, 1999. 

42. Ewer, M. S., Martin, F. J., Henderson, C., Shapiro, C. L., Benjamin, R. S., and Gabizon, A. A., Cardiac 

safety of liposomal anthracyclines, Semin. Oncol., 31(6 Suppl. 13), 161–181, 2004. 

43. Allen, T. M. and Martin, F. J., Advantages of liposomal delivery systems for anthracyclines, Semin. 

Oncol., 31(6 Suppl. 13), 5–15, 2004. 

44. Vail, D. M., Amantea, M. A., Colbern, G. T., Martin, F. J., Hilger, R. A., and Working, P. K., 

Pegylated liposomal doxorubicin: Proof of principle using preclinical animal models and pharmacokinetic 

studies, Semin. Oncol., 31(6 Suppl. 13), 16–35, 2004. 

45. Henderson, C. I., Established and emerging roles of liposomal anthracyclines in oncology. Introduction, 

Semin. Oncol., 31(6 Suppl. 13), 1–4, 2004. 

46. Rifkin, R. M., Gregory, S. A., Mohrbacher, A., and Hussein, M. A., PEGylated liposomal doxorubicin, 

vincristine, and dexamethasone provide significant reduction in toxicity compared with 

doxorubicin, vincristine, and dexamethasone in patients with newly diagnosed multiple myeloma, 

Cancer, 106(4), 848–858, 2006. 

47. Gabizon, A., Catane, R., Uziely, B., Kaufman, B., Safra, T., Cohen, R. et al., Prolonged circulation 

time and enhanced accumulation in malignant exudates of doxorubicin encapsulated in polyethyleneglycol 

coated liposomes, Cancer Res., 54(4), 987–992, 1994. 

48. Uziely, B., Jeffers, S., Isacson, R., Kutsch, K., Wei-Tsao, D., Yehoshua, Z. et al., Liposomal doxorubicin: 

Antitumor activity and unique toxicities during two complementary phase I studies, J. Clin. 

Oncol., 13(7), 1777–1785, 1995. 

49. Alberts, D. S., Muggia, F. M., Carmichael, J., Winer, E. P., Jahanzeb, M., Venook, A. P. et al., 

Efficacy and safety of liposomal anthracyclines in phase I/II clinical trials, Semin. Oncol., 31(6 

Suppl. 13), 53–90, 2004. 

50. Huang, S. K., Martin, F. J., Jay, G., Vogel, J., Papahadjopoulos, D., and Friend, D. S., Extravasation 

and transcytosis of liposomes in Kaposi’s sarcoma-like dermal lesions of transgenic mice bearing the 

HIV tat gene, Am. J. Pathol., 143(1), 10–14, 1993. 

51. Krown, S. E., Northfelt, D. W., Osoba, D., and Stewart, J. S., Use of liposomal anthracyclines in 

Kaposi’s sarcoma, Semin. Oncol., 31(6 Suppl. 13), 36–52, 2004. 

52. Safra, T., Muggia, F., Jeffers, S., Tsao-Wei, D. D., Groshen, S., Lyass, O. et al., Pegylated liposomal 

doxorubicin (doxil): Reduced clinical cardiotoxicity in patients reaching or exceeding cumulative 

doses of 500 mg/m 2 , Ann. Oncol., 11(8), 1029–1033, 2000. 

53. Berry, G., Billingham, M., Alderman, E., Richardson, P., Torti, F., Lum, B. et al., The use of cardiac 

biopsy to demonstrate reduced cardiotoxicity in AIDS Kaposi’s sarcoma patients treated with pegylated 

liposomal doxorubicin, Ann. Oncol., 9(7), 711–716, 1998. 

54. Gabizon, A. A., Lyass, O., Berry, G. J., and Wildgust, M., Cardiac safety of pegylated liposomal 

doxorubicin (Doxil/Caelyx) demonstrated by endomyocardial biopsy in patients with advanced 

malignancies, Cancer Invest., 22(5), 663–669, 2004. 


610 


55. O’Brien, M. E., Wigler, N., Inbar, M., Rosso, R., Grischke, E., Santoro, A. et al., Reduced cardiotoxicity 

and comparable efficacy in a phase III trial of pegylated liposomal doxorubicin HCl 

(CAELYX/Doxil) versus conventional doxorubicin for first-line treatment of metastatic breast 

cancer, Ann. Oncol., 15(3), 440–449, 2004. 

56. Kehrer, D. F., Bos, A. M., Verweij, J., Groen, H. J., Loos, W. J., Sparreboom, A. et al., Phase I and 

pharmacologic study of liposomal lurtotecan, NX 211: Urinary excretion predicts hematologic 

toxicity, J. Clin. Oncol., 20(5), 1222–1231, 2002. 

57. Gelmon, K. A., Tolcher, A., Diab, A. R., Bally, M. B., Embree, L., Hudon, N. et al., Phase I study of 

liposomal vincristine, J. Clin. Oncol., 17(2), 697–705, 1999. 

58. Soepenberg, O., Sparreboom, A., de Jonge, M. J., Planting, A. S., de Heus, G., Loos, W. J. et al., Realtime 

pharmacokinetics guiding clinical decisions; phase I study of a weekly schedule of liposome 

encapsulated paclitaxel in patients with solid tumours, Eur. J. Cancer, 40(5), 681–688, 2004. 

59. Boulikas, T., Stathopoulos, G. P., Volakakis, N., and Vougiouka, M., Systemic lipoplatin infusion 

results in preferential tumor uptake in human studies, Anticancer Res., 25(4), 3031–3039, 2005. 

60. Gabizon, A., Tzemach, D., Horowitz, A. T., Shmeeda, H., Yeh, Y., and Zalipsky, S., Reduced toxicity 

and superior therapeutic activity of a Mitomycin C lipid-based prodrug incorporated in Pegylated 

liposomes. Clin. Cancer Res. 12(6), 1913–1920, 2006. 

61. Bandak, S., Goren, D., Horowitz, A., Tzemach, D., and Gabizon, A., Pharmacological studies of 

cisplatin encapsulated in long-circulating liposomes in mouse tumor models, Anticancer Drugs, 

10(10), 911–920, 1999. 

62. Harrington, K. J., Lewanski, C. R., Northcote, A. D., Whittaker, J., Wellbank, H., Vile, R. G. et al., 

Phase I-II study of pegylated liposomal cisplatin (SPI-077) in patients with inoperable head and neck 

cancer, Ann. Oncol., 12(4), 493–496, 2001. 

63. Meerum Terwogt, J. M., Groenewegen, G., Pluim, D., Maliepaard, M., Tibben, M. M., Huisman, A. 

et al., Phase I and pharmacokinetic study of SPI-77, a liposomal encapsulated dosage form of 

cisplatin, Cancer Chemother. Pharmacol., 49(3), 201–210, 2002. 

64. Zamboni, W. C., Gervais, A. C., Egorin, M. J., Schellens, J. H., Zuhowski, E. G., Pluim, D. et al., 

Systemic and tumor disposition of platinum after administration of cisplatin or STEALTH liposomalcisplatin 

formulations (SPI-077 and SPI-077 B103) in a preclinical tumor model of melanoma, 

Cancer Chemother. Pharmacol., 53(4), 329–336, 2004. 

65. Mori, A., Kennel, S. J., van Borssum Waalkes, M., Scherphof, G. L., and Huang, L., Characterization 

of organ-specific immunoliposomes for delivery of 3 0 ,5 0 -O-dipalmitoyl-5-fluoro-2 0 -deoxyuridine in a 

mouse lung metastasis model, Cancer Chemother. Pharmacol., 35, 447–456, 1995. 

66. Asai, T., Kurohane, K., Shuto, S., Awano, H., Matsuda, A., Tsukada, H. et al., Antitumor activity of 

5 0 -O-dipalmitoylphosphatidyl 2 0 -C-cyano-2 0 -deoxy-1-b-D-arabino-pentofuranosylcytosine is 

enhanced by long-circulating liposomalization, Biol. Pharm. Bull., 21(7), 766–771, 1998. 

67. Park, J. W., Hong, K., Kirpotin, D. B., Colbern, G., Shalaby, R., Baselga, J. et al., Anti-HER2 

immunoliposomes: Enhanced efficacy attributable to targeted delivery, Clin. Cancer Res., 8(4), 

1172–1181, 2002. 

68. Pastorino, F., Brignole, C., Marimpietri, D., Cilli, M., Gambini, C., Ribatti, D. et al., Vascular damage 

and anti-angiogenic effects of tumor vessel-targeted liposomal chemotherapy, Cancer Res., 63(21), 

7400–7409, 2003. 

69. Allen, T. M., Sapra, P., and Moase, E., Use of the post-insertion method for the formation of ligandcoupled 

liposomes, Cell Mol. Biol. Lett., 7(3), 889–894, 2002. 

70. Gabizon, A., Horowitz, A. T., Goren, D., Tzemach, D., Shmeeda, H., and Zalipsky, S., In vivo fate of 

folate-targeted polyethylene-glycol liposomes in tumor-bearing mice, Clin. Cancer Res., 9(17), 

6551–6559, 2003. 

71. Harris, L., Batist, G., Belt, R., Rovira, D., Navari, R., Azarnia, N. et al., Liposome-encapsulated 

doxorubicin compared with conventional doxorubicin in a randomized multicenter trial as first-line 

therapy of metastatic breast carcinoma, Cancer, 94(1), 25–36, 2002. 

72. Bomgaars, L., Geyer, J. R., Franklin, J., Dahl, G., Park, J., Winick, N. J. et al., Phase I trial of 

intrathecal liposomal cytarabine in children with neoplastic meningitis, J. Clin. Oncol., 22(19), 

3916–3921, 2004. 



73. Meyers, P. A., Schwartz, C. L., Krailo, M., Kleinerman, E. S., Betcher, D., Bernstein, M. L. et al., 

Osteosarcoma: A randomized, prospective trial of the addition of ifosfamide and/or muramyl tripeptide 

to cisplatin, doxorubicin, and high-dose methotrexate, J. Clin. Oncol., 23(9), 2004–2011, 2005. 

74. Hong, R. L. and Tseng, Y. L., Phase I and pharmacokinetic study of a stable, polyethylene-glycolated 

liposomal doxorubicin in patients with solid tumors: The relation between pharmacokinetic property 

and toxicity, Cancer, 91(9), 1826–1833, 2001. 

75. O’Byrne, K. J., Thomas, A. L., Sharma, R. A., DeCatris, M., Shields, F., Beare, S. et al., A phase I 

dose-escalating study of DaunoXome, liposomal daunorubicin, in metastatic breast cancer, Br. 

J. Cancer, 87(1), 15–20, 2002. 

76. Lu, C., Perez-Soler, R., Piperdi, B., Walsh, G. L., Swisher, S. G., Smythe, W. R. et al., Phase II study 

of a liposome-entrapped cisplatin analog (L-NDDP) administered intrapleurally and pathologic 

response rates in patients with malignant pleural mesothelioma, J. Clin. Oncol., 23(15), 

3495–3501, 2005. 

77. Ozpolat, B., Lopez-Berestein, G., Adamson, P., Fu, C. J., and Williams, A. H., Pharmacokinetics of 

intravenously administered liposomal all-trans-retinoic acid (ATRA) and orally administered ATRA 

in healthy volunteers, J. Pharm. Pharm. Sci., 6(2), 292–301, 2003. 

78. Matsumura, Y., Gotoh, M., Muro, K., Yamada, Y., Shirao, K., Shimada, Y. et al., Phase I and 

pharmacokinetic study of MCC-465, a doxorubicin (DXR) encapsulated in PEG immunoliposome, 

in patients with metastatic stomach cancer, Ann. Oncol., 15(3), 517–525, 2004. 

79. Booser, D. J., Esteva, F. J., Rivera, E., Valero, V., Esparza-Guerra, L., Priebe, W. et al., Phase II study 

of liposomal annamycin in the treatment of doxorubicin-resistant breast cancer, Cancer Chemother. 

Pharmacol., 50(1), 6–8, 2002. 

80. Verschraegen, C. F., Gilbert, B. E., Loyer, E., Huaringa, A., Walsh, G., Newman, R. A. et al., Clinical 

evaluation of the delivery and safety of aerosolized liposomal 9-nitro-20(s)-camptothecin in patients 

with advanced pulmonary malignancies, Clin. Cancer Res., 10(7), 2319–2326, 2004. 

81. Butts, C., Murray, N., Maksymiuk, A., Goss, G., Marshall, E., Soulieres, D. et al., Randomized phase 

IIB trial of BLP25 liposome vaccine in stage IIIB and IV non-small-cell lung cancer, J. Clin. Oncol., 

23(27), 6674–6681, 2005. 

82. Skubitz, K. M. and Anderson, P. M., Inhalational interleukin-2 liposomes for pulmonary metastases: 

A phase I clinical trial, Anticancer Drugs, 11(7), 555–563, 2000. 

83. Rudin, C. M., Marshall, J. L., Huang, C. H., Kindler, H. L., Zhang, C., Kumar, D. et al., Delivery of a 

liposomal c-raf-1 antisense oligonucleotide by weekly bolus dosing in patients with advanced solid 

tumors: A phase I study, Clin. Cancer Res., 10(21), 7244–7251, 2004. 

84. Boulikas, T., Low toxicity and anticancer activity of a novel liposomal cisplatin (lipoplatin) in mouse 

xenografts, Oncol. Rep., 12(1), 3–12, 2004. 

85. Dark, G. G., Calvert, A. H., Grimshaw, R., Poole, C., Swenerton, K., Kaye, S. et al., Randomized trial 

of two intravenous schedules of the topoisomerase I inhibitor liposomal lurtotecan in women with 

relapsed epithelial ovarian cancer: A trial of the national cancer institute of Canada clinical trials 

group, J. Clin. Oncol., 23(9), 1859–1866, 2005. 

86. Yoo, G. H., Hung, M. C., Lopez-Berestein, G., LaFollette, S., Ensley, J. F., Carey, M. et al., Phase I 

trial of intratumoral liposome E1A gene therapy in patients with recurrent breast and head and neck 

cancer, Clin. Cancer Res., 7(5), 1237–1245, 2001. 

87. Colbern, G. T., Dykes, D. J., Engbers, C., Musterer, R., Hiller, A., Pegg, E. et al., Encapsulation of the 

topoisomerase I inhibitor GL147211C in PEGylated (STEALTH) liposomes: Pharmacokinetics and 

antitumor activity in HT29 colon tumor xenografts, Clin. Cancer Res., 4(12), 3077–3082, 1998. 

88. Beutel, G., Glen, H., Schoffski, P., Chick, J., Gill, S., Cassidy, J. et al., Phase I study of OSI-7904L, a 

novel liposomal thymidylate synthase inhibitor in patients with refractory solid tumors, Clin. Cancer 

Res., 11(15), 5487–5495, 2005. 


30 

CONTENTS 

Positively-Charged Liposomes 

for Targeting Tumor Vasculature 


30.1 Introduction ................................................................................................................... 613 

30.2 The Microvascular Endothelium: Highly Accessible Target of Therapeutics ............ 614 

30.3 Arguments for Vascular Targeting Approach over Interstitial Drug Targeting.......... 616 

30.4 Influence of Cationic Charge on Vascular Targeting in Health and in Disease ......... 617 

30.5 Formulation Development: Alternatives to Routine Drug 

Design and Synthesis .................................................................................................... 617 

30.6 Interaction of Cationic Liposomes with Human Endothelial Cells ............................. 618 

30.7 Stealth (PEGylated) Liposome Technology: A Brief Communication ....................... 619 

30.8 PEGylated Cationic Liposomes: Are They Suited for 

Targeting Tumor Vessels? ............................................................................................ 620 

30.9 Cationic Liposomes Improve Anti-Cancer Therapy .................................................... 621 

30.10 Validation and Long-Term Tracking of Tumor Response to Therapy ....................... 623 

30.11 Concluding Remarks..................................................................................................... 623 

References .................................................................................................................................. 624 


In order to abolish the function of established vessels and/or suppress the formation of neovascular 

networks, various drug delivery vehicles have been used to target experimental therapeutics to the 

tumor vascular surface. Improving site-specific delivery, maximizing tumor response, minimizing 

unwanted side effects, and overcoming tumor heterogeneity are important considerations. 

Clinical strategies are developed in response to unmet clinical needs, but often at the expense of 

understanding how to effectively exploit notable differences between the tumor and host tissues. 

For instance, tumor vessels are generally more tortuous, over-express a variety of active genes, and 

are more permeable to circulating macromolecules (compared to the vasculature of normal tissues). 

On the other hand, relatively high interstitial fluid pressure and long interstitial transport distances 

still hinder effective delivery and deep penetration of drugs in solid tumors. 1 For this reason, 

although the development of interstitial delivery methods represents the focal point of the majority 

of research efforts, alternative treatment options are being developed and evaluated in parallel. 

The justification for parallel research efforts is due to several advantages of endotheliumspecific 

drug targeting over interstitial drug delivery. First, vascular networks in developing 

tumors make tumors vulnerable to treatment regardless of where the tumor is located in vivo. 

Second, the endothelium is vulnerable to the effects of circulating agents; this approach is especially 


613

614 


useful for treating tumors given the fact that a steady flow of oxygen and nutrients (from the blood to the 

interstitial compartment) is essential to sustain active angiogenesis and rapid tumor growth. Third, the 

vast majority of physiological barriers pose little, if any, threat to vascular targeting strategies, and 

finally, endothelial cells are non-mutagenic. This last advantage deals specifically with the inability of 

endothelial cells to acquire drug resistance compared to cancer cells that are inherently more prone to 

mutation and drug resistance. Nonetheless, despite all of these advantages, preservation of normal 

vascular function must be achieved so that treatment is both efficacious and safe. 

The procedures for delivery of drugs to tumor vessels (not the interstitium) have yet to be 

optimized completely for all human organ systems. Whether the aim is to deliver bulky cytotoxic 

drugs or dynamic therapeutic genes, peptides, and proteins, minimizing unintentional injury to 

healthy organ tissues is critical. Until better quality treatments are made available, existing methods 

will have to be applied with a better understanding of their clinical limitations. Successful strategies 

need to satisfy a required set of criteria that include: 

† The drug should not be toxic to patients; 

† Repetitive administration of drug should be well tolerated; 

† The drug should be reasonably accessible to the intended cell target; 

† The drug should be undetectable by the host-immune system; 

† Drug exposure to primary target(s) should yield useful biological product(s); and 

† Biological products should be sufficient to elicit a desired clinical response. 

Although these guidelines represent the gold standard for a number of routine clinical 

procedures and given that some human tumors meagerly respond to standard treatment options 

(yielding inadequate to mediocre responses at best), these routine procedures are simply not 

sufficient for treating all human cancers. 

In general, the endothelium is considered a non-traditional target for FDA approved agents such 

as doxorubicin, vincristine, and paclitaxel; therefore, much of our understanding in connection with 

this approach is incomplete. 

Cytotoxic agents are typically systemically administered at the maximum tolerated dose 

(MTD). The goal has long been to deliver huge therapeutic doses to eradicate millions of viable 

cancer cells hidden deep inside tissues. The drugs next pass through a complicated tumor interstitial 

matrix to prevent the metastatic process from reaching the point of no return. Today, clinical 

success has been achieved with the use of some interstitial drug targeting practices. The question 

currently asked is: Can the use of popular MTD strategies be replaced with an approach that focuses 

on selective delivery of drugs to the tumor vascular compartment rather than to the interstitial tumor 

compartment? Moreover, will this clinical practice result in a more effective clinical response that 

is both efficacious and safe? To address these question and related concerns, the structural and 

functional role of the endothelium should be examined; identification of accessible targets along the 

endothelium is a good start. 

30.2 THE MICROVASCULAR ENDOTHELIUM: HIGHLY ACCESSIBLE TARGET 

OF THERAPEUTICS 

The endothelium is a structural barrier separating the intravascular compartment from the tissue 

space. Because of wide variability in anatomical structure, the vascular compartment is further 

grouped into three different subcategories. The subcategories are continuous, fenestrated, or 

discontinuous endothelia. 

Continuous endothelia are the most common type and are found most frequently in blood 

vessels lining chambers of the heart, walls of capillaries and arterioles in skeletal, skin, cardiac 

muscle, and connective tissue. They are well known for their relatively tight cellular junctions. 2,3 

This particular category of vessels is also critical in the regulation and rapid exchange of ions and 


Positively-Charged Liposomes for Targeting Tumor Vasculature 615 

solutes. 2,3 Plasmalemmal vesicles involved in endothelial transport are abundant in myocardial 

endothelia but are far less frequently observed in capillaries of the brain. 4 The actual number of 

plasmalemmal vesicles existing along the continuous endothelium is heterogeneous, varying as a 

function of organ and tissue environment. These vesicular structures are also highly sensitive to 

charge characteristics, favoring associations with anionic over cationic molecules. 5 

Fenestrated vessels are normally found in vessels of organs that secrete (or excrete) biological 

fluids as in the gastrointestinal mucosa and in the glomerular capillaries of the kidneys. Fenestrae 

are usually between 50 and 80 nm in size, and they appear either as individual gap openings in the 

wall of functional vessels or as clusters. Similar to plasmalemmal vesicles, their frequency of 

occurrence along vessels depends on organ type and microenvironment. 

Often, two or more capillaries may join to form post-capillary venules. These newly formed 

networks are composed of a single lining of endothelial cells with a basement membrane with no 

smooth muscle cell attachment. These vessels are heavily involved in the exchange of molecules, 

and they are preferential sites of plasma extravasation as a result of the actions of vasoactive and 

humoral factors. Fenestrated vessels possess a negatively charged surface density as a result of high 

heparin sulfate proteoglycan content, and unlike plasmalemmal vesicles of continuous endothelia, 

they favor interaction with cationic over anionic molecules. 6 

Discontinuous endothelia are found primarily in the liver, spleen, and bone marrow organs. 2 In 

the liver sinusoids, the endothelia are not continuous, and they possess an average fenestrae size 

between 100 and 150 nm in diameter with the size of fenestrate often changing in response to local 

mediators. These changes include, but are not limited to, response to luminal pressures and potent 

vasodilators such as histamine and bradykinin. 7–10 An investigation into the size of vascular pore 

openings in animal tumor models revealed gap openings that were significantly larger than those 

observed along vessels in normal tissues, around 4 mm (4000 nm) in at least one tumor type but 

normally falling within the range of 0.4–0.6 mm (400–600 nm) for many tumor types. 11,12 Nonetheless, 

the evidence is overwhelmingly in favor of the development of tumor-targeted delivery of 

therapeutic carrier molecules that are small enough to enter through tumor vascular pores without 

passing through openings in normal healthy organ tissues. 

The endothelium is responsible for synthesizing a variety of molecules, regulating endothelial cell 

migration, proliferation, blood vessel maturation, and function. It has been shown to synthesize 

vascular growth factors, nitric oxide, collagen IV, laminin, glycosaminoglycans, and proteoglycans 

to highlight several proven functions. 10,13–19 The physical barrier organizes very rapidly to form 

monolayers, reassembles to form vascular tubes, 20 and can change specific inter- and intra-cellular 

signaling patterns to meet highly specialized needs of the host. Additionally, endogenous and exogenous 

mediators of immune and inflammatory response regulate specialized functions at the surface of the 

endothelium. The endothelium can be considered an effective mediator of organ homeostasis. 10,21 

In many ways, the vascular networks found in solid tumors poorly resemble the more regular, 

well-defined vascular structure observed in disease-free tissues. Tumors, for example, have a highly 

chaotic arrangement of vessels compared to vessels in normal tissues. Tumor vessels also have an 

over abundance of anionic phospholipids in addition to a number of other negatively charged 

functional groups. 22–25 In view of the negatively charged molecules, glycosaminoglycans carry 

out important functions in the metastatic disease process and, much like phospholipids, can serve as 

useful targets of cationic drug carrier molecules. Evaluation of altered proteoglycan expression in 

human breast tissue revealed a total proteoglycan content that was significantly increased in 

comparison to that in healthy tissues. 26 Proteoglycans isolated from malignant breast tissue have 

been shown to stimulate endothelial cell proliferation, and the total glycoprotein content in tumors 

is produced by many cell types, including cancer and tumor endothelial cells alike. 

The vascular networks of tumors have an increased permeability for macromolecules and a 

higher proliferation rate of endothelial cells compared to vessels in quiescent tissues. 27,28 An 

estimated 30- to 40-fold increase in the growth rate of endothelial cells lining vessels in tumors 

compared to that in normal tissues has been demonstrated. 27 


616 


Direct access to intravenously administered agents, rapid proliferation rate of endothelial cells, 

and over-expression of negatively charged functional groups along vessels are potentially exploitable 

features of tumors. Endothelium-specific cationic liposomes made more specific through 

conjugation of antibodies can offer tremendous therapeutic gains. 

30.3 ARGUMENTS FOR VASCULAR TARGETING APPROACH OVER 

INTERSTITIAL DRUG TARGETING 

Physiological barriers that prevent effective interstitial delivery and drug transport 29 will likely pose 

little threat to vascular targeting strategies given that blood vessels have better access to intravenously 

administered agents. Damage to only a few tumor vessels can result in the death of literally thousands 

of malignant cells. 30 Limited drug diffusion to hard-to-reach interstitial areas in tumors will ultimately 

spare clusters of well-differentiated cancer cells from treatment, but direct targeting of cancer cells is 

not as important for vascular targeting strategies. As a therapeutic approach, vascular targeting is 

unaffected by convection and diffusion, and it does not require the presence of highly perfused vessels 

in all tumor regions to be effective. It can be argued that irregular, slow blood flow velocities play a 

critical role in facilitating interaction of drugs with endothelial cells. 31 Additionally, poorly perfused 

tumor vessels permit limited access of drugs to cancer cells. Additionally, endothelial cells are nonmalignant 

and are far less likely than cancer cells to acquire the mutations that safeguard them from 

treatment. Finally, should a drug intended to target functional tumor vessels gain direct access to a 

population of cancer cells in the process, additional gains for therapy could result. The potential for 

vascular targeting substances to come in direct contact with cancer cells can be observed in mosaic 

vessels where tumor cells directly interact with the lumen of vessels and flowing blood and assist in 

the spread of metastatic disease. 32 This is apparent with large tumor vascular pore openings where 

drug agents have considerably greater access to malignant cell populations. 11 

Proof-of-principal: A significant amount of experimental evidence can now be used to 

support vascular targeting as a targeted approach against cancer. To illustrate, tumor-bearing 

mice treated with doxorubicin–RGD-4C conjugate (tumor-homing peptide), outlived control 

mice that died from widespread disease. 33 Mice treated every 3 weeks for a total of 12 weeks 

with doxorubicin-RGD-4C lived 6 months longer compared to doxorubicin-treated mice. 33 An 

in vitro study showed success with immunoliposome-targeted delivery to ICAM-1-expressing 

cells in the bronchial epithelium and endothelium. 34 In this study, the extent of ICAM-1 

expression was suggested to influence the degree of immunoliposome binding to endothelial 

cells. The immunotargeting of liposomes to activated vascular endothelial cells of the cardiovascular 

system has also been investigated. 35 In this study, murine antibody mAB H18/7 was 

conjugated to doxorubicin-loaded liposomes, and the extracellular binding domain of E-selectin 

was targeted to a significant extent over nonactivated human endothelial cells. 35 Thorpe and 

colleagues characterized TEC-11 antibody and have demonstrated that the anti-class-ricin 

A-chain immunotoxin produced widespread infarction and tumor regression in treated mice 

compared to untreated controls. 36 While these groups and others work to demonstrate proofof-principal, 

other laboratories have had success with identifying accessible tumor vessel 

targets. One example of a potentially useful clinical target was studied by Retting and 

colleagues. They confirmed that the tumor endothelial antigen endosialin (cell glycoprotein) 

is expressed in tumors and is not expressed in normal healthy tissues. 37 This early finding 

suggests that, under optimized conditions, the glycoprotein could be used to launch a sitespecific, 

immunological attack on tumor vessels. Multiple lines of evidence from the literature 

support vascular targeting as an attractive therapeutic approach against cancer disease. 



30.4 INFLUENCE OF CATIONIC CHARGE ON VASCULAR TARGETING 

IN HEALTH AND IN DISEASE 

To optimally target drugs to the endothelium, a basic working knowledge of why select carriers 

accumulate at this site is essential. Extraordinary insights into normal vessel function and pathogenesis 

now provide some interesting clues. For example, the vessel lumen of capillaries is 

predominately lined with anionic sites. 22–25 These regions are important in the development and 

maintenance of normal and abnormal tissues where the latter is observed in inflammation, thrombosis, 

and in tumors. 24,25,38 Anionic sites (owing to layers of glycoproteins) mediate microvascular 

permeability and trans-endothelial cell function. 3,6,23–25,30,38 Endothelial cell membrane-associated 

proteoglycans regulate delivery and uptake of macromolecules. This has been demonstrated in 

sulfated proteoglycan deficient HeLa cells where cation-mediated delivery of plasmid DNA 

resulted in a 69% reduction in luciferase expression. 39 The distribution of negatively charged 

sites is non-uniform, 25,31 varying among vessel types (i.e., arteries, capillaries, or capillary 

venules) and between young and relatively old mice. 4 Although negatively charged components 

of mammalian cells are sialoglycoproteins and proteoglycans, some vascular domains (potentially 

involved in endocytosis and/or trancytosis) are devoid of these functional groups. These specialized 

areas interact to a lesser extent with cationized molecules, and it has been suggested that they 

represent regions of anionic molecular transport across capillary networks. 25 The absence of 

accessible anionic sites can be observed on the vascular surface, specifically on the surface of 

plasmalemmal vesicles where they readily assist in the transport of circulating macromolecules. 4,6 

Experiments in solid tumors support the notion that success with nanoparticle-assisted delivery 

of drugs to tumor vessels does not require that drug carrier molecules possess a cationic charge for 

effective delivery and transport. In fact, when anionic sites were neutralized with the polycation 

protamine, the blood–brain barrier (BBB) became significantly less guarded against the transport of 

albumin and inulin; under normal conditions, neither can pass the BBB. 40 Additionally, whereas 

electron microscopic studies revealed that positively charged electron-dense label cationic ferritin 

(CF) did not label plasmalemmal vesicles of the mouse endothelium, native anionic and neutralized 

ferritin molecules had considerably greater access to these sites. 4 Another study showed that 

following a cerebral insult of a rat brain tumor model, the vasculature was more permeable to 

horse radish peroxidase, but more importantly, showed a fewer number of negatively charged sites. 

The mechanism(s) by which anionic surface charge is altered in disease is not completely understood, 

40–42 but studies continue to support structural and morphological differentiation of the 

microvascular endothelium in disease tissues that undoubtedly influences overall vascular 

targeting efficiency. 

The endothelium is structurally and functionally unique, and several studies confirm regionspecific, 

tumor vascular sensitivity to effects of molecular charge. 3,6,18,25,40,43,44 The fact that the 

endothelium can bind anionic, neutral, and net positively charged molecules supports these observations. 

The strategies that best exploit the expression of negatively charged functional groups in 

combination with additional tumor vascular features will achieve the greatest therapeutic success. 

30.5 FORMULATION DEVELOPMENT: ALTERNATIVES TO ROUTINE DRUG 

DESIGN AND SYNTHESIS 

The development of novel therapeutics is an expensive and labor-intensive process. There is also no 

guarantee that a more useful therapeutic agent will result from the time and money invested. The use of 

combinatorial chemistry, high throughput screening, and drug synthesis has not met all clinical expectations. 

As a result, serious issues still remain with existing treatment approaches and with a search for a 

magic bullet. 45 For this reason, strategies that improve tumor endothelial site-specific delivery are 

being developed in parallel with efforts involving routine drug design and synthesis. 


618 

In many instances, the use of nanotechnology represents a wise alternative to routine drug 

synthesis. This technology creates new opportunities to improve drug access to primary targets and 

to exploit additional tumor targets as well. Long circulating liposomes, for example, increase 

deposition of drugs in the perivascular space of tumors. 46,47 When the optimal lipid ratios have 

been considered, the vasculature may be targeted more selectively compared to free drug. 

On the basis of the experimental evidence, a non-toxic phospholipid-based delivery vehicle is a 

rational choice for targeting many therapeutic agents to tumors. Liposomes are well suited for this 

purpose. These microparticulates spontaneously form and are generally easier to prepare compared 

to viral mediated systems. 

Liposomes are relatively non-immunogenic; their large-scale manufacturing does not depend 

on the culturing of living cells, and they are preferred for a variety of practical considerations as 

well. Liposomes have been shown to increase systemic circulation, reduce drug side effects, 

enhance cellular uptake, and increase the duration of drug exposure and tumor response to 

therapy. 16–18 Moreover, the physicochemical properties of liposomes can be manipulated to 

address physiological considerations such as the effects of size, molecular charge, membrane 

rigidity, and stability. 48–53 

Developing a target specific, long circulating, stable drug carrier is only half the battle. Ideally, 

liposomes should also be triggered to release drugs under tumor sensitive conditions. The effect of 

molecular charge, over-expression of vascular growth factors, sensitivity to pH, and relatively large 

vascular pore openings are a few physiological features currently used to distinguish tumors from 

healthy organ tissues. 12,25,31,38,46,53–57 

30.6 INTERACTION OF CATIONIC LIPOSOMES WITH HUMAN 

ENDOTHELIAL CELLS 


One of the first synthetic cationic lipids, developed by Felgner and colleagues in 1987, 58,59 was 

initially used to improve existing DNA delivery methods. A 1:1 w/w mixture of the new cationic 

lipid DOTMA (N-[1-(2,3-dioleyloxy) propyl]-N,N,N-trimethylammonium chloride) was combined 

with the helper lipid DOPE (dioleoylphosphatidylethanolamine). This formulation (akaw- 

Lipofectin) was found to be 5–100 times more effective than DNA transfection performed with 

other in vitro methods. 59 A typical eukaryotic cell membrane has a highly negative surface charge 

density because of the multitude of membrane-bound carbohydrate groups oriented on its surface. 

Endothelial cells can be targeted on the basis of this surface charge density, and several 

mechanisms of action have already been proposed for successful delivery of DNA into cells by 

positively charged liposomes. 60–63 

Studies involving the transfection of human endothelial cells suggest that methods be tailored 

to specifications of endothelial cells, not more traditional cell types. High transfection efficiency 

with the use of cationic liposomes is possible when expression is under the control of a strong 

promoter. 64 Endothelial cells, like most mammalian cells, depend on an optimal DNA/cationic 

liposome ratio for high efficiency of endothelial cell trafficking and expression of foreign genes. 

Moreover, compared to other mammalian cell lines, primary endothelial cells are generally more 

sensitive to constituents of growth medium and passage number. 65 No specific cationic lipid, to 

date, is accepted for use with all mammalian cell types. Cationic liposomes may represent a suitable 

alternative to non-viral delivery systems; however, transfection of human endothelial cells by this 

approach should be optimized to produce the desired therapeutic effect. 64 

Various experimental and clinical applications for cationic lipids have been reported, ranging 

from optimizing methods for enhanced gene transfer in vitro to treating clinical abnormalities 

contributing to neoplastic transformations in man. 66 Using cationic lipids to regulate endothelial 

cell function in vitro or in treatment of disease in highly developed organ systems has been 

explored. To illustrate, the application of several cell line types have been investigated 



FIGURE 30.1 (See color insert following page 522.) Cationic liposomes are taken up by human endothelial 

cells. The figure shows a DIC and fluorescence merged image of PCLs associated with human dermal 

endothelial cells. Perinuclear localization is observed. Captured image is representative of PCLs taken up 

by other primary endothelial cells derived from various organ tissues. Cells were seeded at 5!10 5 cells/well on 

cover slips in six well plates. Images were captured 24 h after exposure of cells to PCLs. 

(Figure 30.1). Cationic liposome-mediated transfection has been carried out under a broad set of 

experimental conditions, involving the use of numerous in vitro models of the endothelium in 

human organ tissues. These in vitro models of the intravascular compartment include the eye, 67 

lung, 68,69 heart, 61,70 kidney, 71 and vascular smooth muscle. 72,73 

Site-specific delivery to endothelial cells can be accomplished by coating cationic lipsoemes 

with peptides and monoclonal antibodies. 65,74 Furthermore, these positively charged vehicles 

enhanced adeno-associated virus (AAV) uptake by human endothelial and smooth muscle cells. 

The authors report a highly efficient way to introduce foreign DNA into cells given that modified 

cationic liposomes (adenosomes) enhanced vector expression by a factor of 10 over adenoviral 

proteins delivered without liposomes. 75 

Although antibodies have been used to assist gene transfer strategies, cationic liposomes can be 

optimized to deliver genes to endothelial cells without the use of antibodies. For instance, when 

cultured bovine corneal endothelial cells (BCECs) were exposed to pUT651 (a plasmid) alone, 

significant expression was observed and without associated toxicities. 67 In HUVEC cells, reporter 

gene expression was evident in the absence of select antibodies. In this study, transfection efficiency 

varied as a function of cell passage number, and a linear reduction in luciferase expression was 

observed with successive passage numbers. 65 Vascular smooth muscle cells were shown to vary in 

degree of sensitivity to transfection reagents mediated by cationic lipids, reporting a preference for 

restenotic lesions compared to normal internal thoracic arteries and atherosclerotic plaques. 73 

30.7 STEALTH (PEGYLATED) LIPOSOME TECHNOLOGY: A BRIEF 

COMMUNICATION 

Rapid elimination of cationic liposomes from circulation by opsonins, lipoproteins, and cell-surface 

receptors pose an additional threat to effective delivery of drugs to solid tumors. 45,76–80 This 

particular barrier is different from physiological barriers as it results not from tumors or the 

influences from the surrounding tissue environment, but it occurs when conventional (relatively 

short circulating) liposomal therapeutics come in direct contact with circulating proteins in blood. 

Stealth liposome technology is one approach used to limit the interaction of conventional 

liposomes with proteins in blood. The technology involves the coupling of high molecular 


620 

weight polymers to the liposome surface. 47,50 Surface-bound polymers [i.e., poly(ethylene glycol) 

(PEG)] provide adequate protection of liposomes from protein molecules that can unfavorably alter 

their tissue distribution profile. The specific mechanisms underlying delivery mediated by these 

polymers are unclear, but when PEG is included as a component of liposome preparations, a 

physical zone is formed around the liposome perimeter. This zone of exclusion reduces liposome–protein 

interactions as a result of long (PEG) polymer chain constrictions. 47,50,80 

From a therapeutic point-of-view, in addition to favorable kinetics, PEGylated liposomes are 

usually more active against solid tumors. 47,79 Several lines of evidence support the findings that 

PEGylated liposomes are generally more stable in blood with only minor leakage of their contents 

over time compared to conventional liposomal therapeutics. 78 

PEG can also be used to prepare liposomes with significantly higher drug retention under 

normal physiological pH, or it can be designed to trigger drug release in a low pH environment. 

In addition to tumor uptake, other organs of relatively high accumulation are the liver, kidney, skin, 

the female reproductive, and gastrointestinal tract. 77,78 

The extent of modification and molecular mechanics of Stealth liposome technology and 

exactly how PEG delays reticuloendothelial system (RES) entrapment is not completely understood. 

However, its benefit(s) in systemic circulation and that the fraction of PEG included in 

Stealth liposome preparations should reflect preclinical and clinical goals are understood. 

Stealth technology might well represent one of the most significant advances in the field of 

liposome nanotechnology. PEG has provided researchers with an enormous opportunity to exploit 

the large openings of tumor vessels by enhancing delivery of therapeutic substances through them 

as a result of the advantage of prolong circulation. In a study performed by Yuan et al., Stealth 

liposomes accumulated in the perivascular space in tumors 48 h post-administration. 46 This is the 

region close to tumor vessels but not directly associated with them. The technology is, however, 

not sufficient to promote vascular-specific delivery; targeting specific physiological feature of 

vessels (i.e., proteoglycans, endothelial receptors, etc.) in addition to prolonged circulation is 

also required. 

30.8 PEGYLATED CATIONIC LIPOSOMES: ARE THEY SUITED 

FOR TARGETING TUMOR VESSELS? 


The notion of targeting accessible anionic sites along tumor vessels combined with technologies 

that limit interactions of nanodelivery systems with biological molecules in vivo is promising. 

PEGylated cationic liposomes (PCLs) have been shown to improve not only oligonucleotide 

(ODN) loading and delivery, but also the apparent drug solubility profile of various therapeutic 

molecules as well. In vitro studies conducted with human plasma show that PCLs do not form large 

molecular aggregates in blood 81,82 or in buffered solutions. 31 These carriers retain most of their 

physicochemical properties, 31,75,81 and they prolong circulation half-life even in the presence of a 

net positively charged surface. 

Protecting cationic liposomes from serum nucleases enhanced immunostimulatory activity of 

CpG oligonucleotides motifs compared to free ODN in in vitro and in mouse models. 82 In another 

study, the anti-HER-2 F (ab 0 ) fragment was coupled to the distal end of PEG chains of cationic 

liposomes; cellular uptake and nuclear localization increased considerably in cultured SK-BR-3 

cells (a human breast cancer cell line) shown to over-expresses HER-2 oncoprotein. 81 

Intravital microscopy has been used to investigate the interactions of PCLs with tumor vessels 

in vivo. 31 When PCLs were intravenously injected into tumor bearing (LS174T-human colon 

adenocarcinoma) mice, analysis of spatial distribution of liposomes 24 h post-injection revealed 

significant vascular targeting in tumors (w27.5% coverage) compared to vessels in normal tissues 

(w3.6% coverage). 31 Moreover, the extent of vascular areas targeted did not vary as a function of 

tumor type or anatomical location. 31 Several lines of evidence suggest that PEG coating does not 



inhibit cationic liposomes from binding to negatively charged surfaces. In addition to experimental 

data supporting this hypothesis, the values of zeta potential (electric potential across a double 

membrane surface) confirm that a sufficient electrostatic membrane potential still exists in the 

presence of PEG. 31,52,83 

Vascular targeting can be influenced on the basis of charge content of drug carrier molecules; 

associations with tumor vessels favor cationic liposomes possessing relatively high cationic charge 

content. 31,38,63 The abilities to circulate for extended periods in blood and associate with tumor 

vessels are highly desirable features of an efficient therapeutic approach. For this reason, both the 

inclusion of PEG and cationic charge content in cationic liposomes are currently being investigated 

in combination. There are clinical circumstances when relatively long circulating drug carrier 

molecules (O24 h) accumulating along tumor vessels would be more effective than carriers that 

are rapidly eliminated from circulation (!0.5 h). For instance, many of the popular chemotherapeutic 

agents are dependent on phase of the cell cycle; therefore, rapidly dividing tumor endothelial 

cells could be targeted more effectively with carriers that maintained adequate levels of drug in 

blood. Such a delivery system would serve less value for agents that can exert drug effects outside 

the influence of the cell cycle (i.e., membrane targeting agents). 

30.9 CATIONIC LIPOSOMES IMPROVE ANTI-CANCER THERAPY 

The use of cationic liposomes to deliver experimental therapeutics to solid tumors is a promising 

alternative to interstitial targeting approaches. However, formulations should be optimized and prescreened 

in preclinical models to establish clinical relevance. Given the fact that the architecture of 

tumor vessels varies significantly in terms of microvascular type, density, vascular surface area, and 

angiogenic potential, it is understandable why the development of a single therapeutic approach 

against cancer has yet to be achieved. 

When developing cationic liposomal therapeutics, it will be helpful to note that each solid 

tumor possesses distinctly different organization and specialized structural features. The highly 

unpredictable nature of tumors further complicates the task of pairing formulations to specific 

disease states. A common feature among all cationic liposome preparations used in drug delivery 

is that they must be optimized for site-specific delivery to tumor endothelia and that their intracellular 

fate should correlate with the drug’s mechanism of action (i.e., cationic liposomes should 

enhance delivery of DNA targeting agents to DNA, and not to irrelevant intracellular targets). This 

process must be confirmed. 

Cationic liposomes can retain drug agents at the tumor vascular site (Figure 30.2), and facilitate 

interaction of liposomes with subcellular targets prior to releasing their payload. They can also be 

used to target non-intracellular targets as well as cell-membrane bound molecules other than 

proteoglycans. This is promising given that many anti-angiogenic agents have been confirmed to 

exert their action by each mechanism. For instance, SU5416, an inhibitor of tyrosine kinase activity 

of vascular endothelial growth factor (VEGF), requires direct access to a specific endothelial cell 

membrane-associated receptor (Flk-1/KDR) in order to suppress neovascular growth of tumors. 84,85 

Although SU5416 is suggested to exert long-lasting effects on VEGF phosphorylation and function, 

cationic liposome-assisted drug delivery could enhance interactions with specific endothelial 

cell targets. 

Effective anti-angiogenic therapy requires the continuous presence of drugs in circulation. 85,86 

The inclusion of PEG in cationic liposome preparations can extend circulation half-life of SU5416 

compared to SU5416 alone; the goal here is to enhance the duration of drug (SU5416) exposure 

with tumor target. 

Most anti-angiogenic agents suppress neovascularization by blocking activators of angiogenesis 

and endothelial cell-specific signals. There exist clinical situations where targeting nontumoral 

tissues might be equally beneficial. Although the majority of PCLs are recovered by the 


622 


FIGURE 30.2 (See color insert following page 522.) Targeting tumor vessels with cationic liposomes. The 

image shows vascular targeting of LS174T (human colon adenocarcinoma) with cationic liposomes and was 

captured by two-photon microscopy 250 mm deep within the tumor. The intrinsic fluorescent property of 

doxorubicin was used to probe the location of the drug within the tumor. When doxorubicin was loaded in 

PCLs and administered via tail vein in tumor bearing mice, the drug was shown to be delivered to tumor 

vessels. Green represents tumor vessels, and orange dots represent doxorubicin. Mag. BarZ50 mm. 

liver post-intraveneous 31 and intrasplenic 87 administration, once these carrier molecules have been 

loaded with chemotherapeutic agents, they can diminish primary tumor growth in this anatomical 

location without significantly altering its primary function. To illustrate, poly(I):poly(C)–cationic 

liposome complex was used to treat metastatic (pancreatic and colorectal) carcinoma cell growth in 

the liver, and the result was significant delay in tumor growth compared to the administration of 

poly(I):poly(C), and adriamycin alone. 87 

In a study using PCLs to deliver the chemotherapeutic agent cisplatin (CDDP), CDDP-loaded 

PCLs suppressed primary tumor growth and distant liver metastasis as a result of tumor and liverselective 

uptake. 88 The authors note that the high chondroitin sulfate content of cancer cells was the 

primary target of PCLs. In addition to this mechanism, another possibility is that drug loaded-PCLs 

deliver drugs like CDDP to tumor vessels and that cancer cells are not the primary, but the 

secondary, target of drug-loaded PCLs. Depending on the overall clinical objectives, liver accumulation 

could result in some unwanted complications; for wide-spread general applicability, high 

tumor selectivity at the expense of the liver is desired. To minimize uptake of cationic carriers at 

this anatomical site, additional specialized features of tumor vessels may have to be taken 

into account. 

In addition to targeting tumor vascular targets on the basis of negatively charged functional 

groups, cationic drug carriers can also be used to enhance the interaction of anti-angiogenic agents 

with integrins. Integrins are involved in cell invasion and metastasis and in signaling processes that 

regulate the actions of each. 89,90 Some integrins are preferentially expressed in angiogenic endothelium 

like integrin avb3. 89 Researchers have since introduced a novel cationic delivery vehicle in 



order to exploit this endothelial cell receptor. 91 In as study carried out by Hood and colleagues, a 

cationic lipid based nanoparticle (NP) was coupled to avb3-targeting ligand and intravenously 

injected in tumor bearing mice; the complex (avb3–NP) selectively targeted the vasculature and 

induced tumor regression. 91 Furthermore, sustained regression was achieved without significant 

accumulation in, and damage to, the liver. This finding is encouraging. It supports the use of 

positively charged drug carriers with highly specialized molecules to improve vascular targeting 

at all stages of tumor development. 

The knowledge that cationic liposomes deliver drugs to solid tumors as well as the degree to 

which each tumor compartment is affected by treatment is essential to future advances in nanotechnology. 

Unfortunately, the overwhelming majority of published works document overall tumor 

response, animal survival, and drug effects involving the interstitial matrix, yet hidden 

mechanism(s) underlying tumor vascular damage in connection with treatment is often not a part 

of routine investigations. 

No intravenously administered substances (drug alone or nanoparticle-assisted) can gain access 

to the interstitial matrix without first passing through the vascular compartment, so drug effects to 

this tumor compartment must be explored. Cationic liposomes possessing a high cationic lipid 

content (w50 mol%) target the endothelium several-fold over the interstitial environment. 31,38 

Growth inhibition and animal survival studies have proven to be extremely helpful in establishing 

safety conditions for drug administration and predicting drug efficacy in humans. Nonetheless, the 

specific mechanisms by which cationic liposomes target vessels and alter tumor growth kinetics 

must be investigated. Studies should probe interstitial and vascular compartments without 

experimental bias. 

30.10 VALIDATION AND LONG-TERM TRACKING OF TUMOR RESPONSE 

TO THERAPY 

It has been established that there are unique molecular and physiological features of tumors that can 

be used to develop relatively disease-specific treatments. The expression of tumor-specific genes 

has been used to monitor progression of disease in response to these treatments and others. For 

instance, VEGF is a potent angiogenic cytokine that is critically linked to angiogenesis; it is so vital 

to this process that inhibitors are rapidly being developed to suppress VEGF’s tumor angiogenesis 

stimulatory activity. 84,92–94 Normally, when select inhibitors of VEGF bind Flt-1 (VEGF target 

site) on endothelial cells, proliferation is inhibited among other critical steps involved in tumor 

angiogenesis. 20,30,84 VEGF expression is a highly regulated process in healthy tissue; elevated 

expression is observed only under special circumstances, but constitutive expression is observed 

in tumors that aggressively seek to recruit a new blood supply. Given the fact that the process of 

developing new blood vessels is fundamentally important to maintaining an aggressive tumor 

phenotype, VEGF and a variety of other growth factors are often used as molecular and 

pharmacological endpoints. 

Site-specific delivery to tumor vessels and angiogenesis has demonstrated promise in preclinical 

models, warranting follow-up investigations in clinical settings. Evaluating tumor volume, 

animal survival and vascular density are useful indicators of tumor response. However, monitoring 

gene expression in response to therapy can provide deeper insights into the specific mechanisms 

underlying the disease process. 


Systemic administration of drugs is a frequently applied strategy used to control local tumor growth 

and metastasis. The issue of targeting these agents more specifically to tumors while reducing 


624 

uptake by healthy organ tissues is a formidable challenge. Cationic liposomes have been used to 

mediate efficient transfection of a variety of different mammalian cell types. The loading of cationic 

liposomes with chemotherapeutic agents has even greater potential. The relatively new class of 

liposomes possesses high cationic charge content (and membrane surface charge potential) similar 

to those used for transfection studies, but they deliver their drug payload to tumor vessels. The 

development of new synthetic cationic lipids like N-[1-(2,3-dioleyloxy)propyl]-N, N, N-trimethylammonium 

chloride(DOTMA) and N-[1-(2,3-dioleyloxy)propyl]-N, N, N-trimethylammonium 

chloride(DOTAP) has helped make this field a reality. 

A deeper understanding of how tumors develop relative to the expression of vascular receptors 

and other targets will create new opportunities for nanotechnologists. These opportunities extend to 

include more effective treatments at various stages of tumor development and planned attacks 

against cancer. 

REFERENCES 


1. Jain, R. K., Vascular and interstitial barriers to delivery of therapeutic agents in tumors, Cancer 

Metastasis Rev., 9, 253–266, 1990. 

2. Simionescu, M., Simionescu, N., and Palade, G. E., Morphometric data on the endothelium of blood 

capillaries, J. Cell Biol., 60, 128–152, 1974. 

3. Simionescu, M., Simionescu, N., and Palade, G. E., Structural basis of permeability in sequential 

segments of the microvasculature. I. Bipolar microvascular fields in the diaphragm, Microvasc. Res., 

15, 1–16, 1978. 

4. Simionescu, M., Simionescu, N., Santoro, F., and Palade, G. E., Differentiated microdomains of the 

luminal plasmalemma of murine muscle capillaries: Segmental variations in young and old animals, 

J. Cell Biol., 100, 1396–1407, 1985. 

5. Ghinea, N. and Simionescu, N., Anionized and cationized hemeundecapeptides as probes for cell 

surface charge and permeability studies: Differentiated labeling of endothelial plasmalemmal 

vesicles, J. Cell Biol., 100, 606–612, 1985. 

6. Simionescu, M., Simionescu, N., and Palade, G. E., Differentiated microdomains on the luminal 

surface of the capillary endothelium. I. partial characterization of their anionic sites, J. Cell Biol., 

90, 614–621, 1981. 

7. Movat, H. Z., The role of histamine and other mediators in microvascular changes in acute inflammation, 

Can. J. Physiol. Pharmacol., 65, 451–457, 1987. 

8. Mylecharane, E. J., Mechanisms involved in serotonin-induced vasodilation, Blood Vessels, 27, 

116–126, 1990. 

9. Vercellotti, G. M. and Tolins, J. P., Endothelial activation and the kidney: Vasomediator modulation 

and antioxidant strategies, Am. J. Kidney Dis., 21, 331–343, 1993. 

10. van Hinsbergh, V. W., The endothelium: Vascular control of haemostasis, Atherosclerosis, 95, 

198–201, 2001. 


microenvironment, Proc. Natl Acad. Sci. USA, 95, 4607–4612, 1998. 


cutoff size, Cancer Res., 55, 3752–3756, 1995. 

13. Hudson, B. G., Reeders, S. T., and Trygvason, K., Type IV collagen: Structure, gene organization and 

role in human diseases, J. Biol. Chem., 268, 26033–26036, 1993. 

14. Gorog, P. and Pearson, J. D., Sialic acid moieties on surface glycoproteins protect endothelial cells 

from proteolytic damage, J. Pathol., 146, 205–212, 1985. 

15. Hallmann, R. et al., Expression of function of laminins in the embryonic and mature vasculature, 

Physiol. Rev., 85, 979–1000, 2005. 

16. Sund, M., Xie, L., and Kalluri, R., The contribution of vascular basement membranes and extracellular 

matrix to the mechanics of tumor angiogenesis, APMIS, 112, 450–462, 2004. 

17. Ruoslahti, E., Structure and biology of proteoglycans, Annu. Rev. Cell Biol., 4, 229–255, 1988. 

18. Ruoslahti, E., Specialization of tumor vasculature, Nat. Rev. Cancer, 2, 83–90, 2002. 



19. Satoh, A., Toida, T., Yoshida, K., Kojima, K., and Matsumoto, I., New role of glycosaminoglycans on 

the plasma membrane proposed by their interaction with phosphatidylcholine, FEBS Lett., 477, 

249–252, 2000. 

20. Folkman, J. and Haudenschild, C., Angiogenesis by capillary endothelial cells in culture, Trans. 

Ophthalmol. Soc. UK, 100, 346–353, 1980. 

21. Michiels, C., Arnould, T., and Remacle, J., Endothelial cell responses to hypoxia: Initiation of a 

cascade of cellular interactions, Biochim. Biophys. Acta, 1497, 1–10, 2000. 

22. Charonis, A. S. and Wissig, S. L., Anionic sites in basement membranes. Differences in their electrostatic 

properties in continuous and fenestrated capillaries, Microvasc. Res., 25, 265–285, 1983. 

23. Roberts, W. G. and Palade, G. E., Neovasculature induced by vascular endothelial growth factor is 

fenestrated, Cancer Res., 57, 765–772, 1997. 

24. Ran, S., Downes, A., and Thorpe, P. E., Increased exposure of anionic phospholipids on the surface of 

tumor blood vessels, Cancer Res., 62, 6132–6140, 2002. 

25. Vincent, S., DePace, D., and Finkelstein, S., Distribution of anionic sites on the capillary endothelium 

in an experimental brain tumor model, Microcirc. Endothleium Lymphatics, 45–67, 1988. 

26. Vijayagopal, P., Figueoroa, J. E., and Levine, E. A., Altered composition and increased endothelial 

cell proliferative activity of proteoglycans isolated from breast carcinoma, J. Surg. Oncol., 68, 

250–254, 1998. 

27. Denekamp, J. and Hobson, B., Endothelial-cell proliferation in experimental tumours, Br. J. Cancer, 

46, 711–720, 1982. 

28. Denekamp, J., Vasculature as a target for tumor therapy, Prog. Appl. Microcirc., 4, 28–38, 1984. 

29. Jain, R. K., The next frontier of molecular medicine: Delivery of therapeutics, Nat. Med., 4, 655–657, 

1998. 

30. Folkman, J., Tumor angiogenesis: Therapeutic implications, N. Engl. J. Med., 285, 1182–1186, 1971. 

31. Campbell, R. B. et al., Cationic charge determines the distribution of liposomes between the vascular 

and extravascular compartments of tumors, Cancer Res., 62, 6831–6836, 2002. 

32. Chang, Y. S. et al., Mosaic blood vessels in tumors: Frequency of cancer cells in contact with flowing 

blood, Proc. Natl Acad. Sci. USA, 97, 14608–14613, 2000. 


vasculature in a mouse model, Science, 279, 377–380, 1998. 

34. Bloemen, P. G. et al., Adhesion molecules: A new target for immunoliposome-mediated drug 

delivery, FEBS Lett., 357, 140–144, 1995. 

35. Spragg, D. D. et al., Immunotargeting of liposomes to activated vascular endothelial cells: A strategy 

for site-specific delivery in the cardiovascular system, Proc. Natl Acad. Sci. USA, 94, 8795–8800, 

1997. 

36. Thorpe, P. E. and Burrows, F. J., Antibody-directed targeting of the vasculature of solid tumors, 

Breast Cancer Res. Treat., 36, 237–251, 1995. 

37. Rettig, W. J. et al., Identification of endosialin, a cancer surface glycoprotein of vascular endothelial 

cells in human cancer, Proc. Natl Acad. Sci. USA, 89, 10832–10836, 1992. 

38. Thurston, G. et al., Cationic liposomes target angiogenic endothelial cells in tumors and chronic 

inflammation in mice, J. Clin. Investig., 101, 1401–1413, 1998. 

39. Mislick, K. A. and Baldeschwieler, J. D., Evidence for the role of proteoglycans in cation-mediated 

gene transfer, Proc. Natl Acad. Sci. USA, 93, 12349–12354, 1996. 

40. Hardebo, J. E. and Kahrstrom, J., Endothelial negative surface charge areas and blood–brain barrier 

function, Acta Physiol. Scand., 125, 495–499, 1985. 

41. Nag, S., Role of the endothelial cytoskeleton in blood–brain-barrier permeability to proteins, Acta 

Neuropathol. (Berl.), 90, 454–460, 1995. 

42. Nag, S., Blood–brain barrier permeability using tracers and immunohistochemistry, Methods Mol. 

Med., 89, 133–144, 2003. 

43. Renkin, E. M., Transport of macromolecules: Pores and other endothelial pathways, Appl. Physiol., 

58, 315–325, 1985. 

44. Renkin, E. M., Multiple pathways of capillary permeability, Circ. Res., 41, 735–743, 1977. 

45. Gabizon, A. A., Stealth liposomes and tumor targeting: One step further in the quest for the magic 

bullet, Clin. Cancer Res., 7, 223–225, 2001. 


626 


46. Yuan, F. et al., Microvascular permeability and interstitial penetration of sterically stabilized (Stealth) 

liposomes in a human tumor xenograft, Cancer Res., 54, 3352–3356, 1994. 

47. Papahadjopoulos, D. et al., Sterically stabilized liposomes: Improvements in pharmacokinetics and 

anti-tumor efficacy, Proc. Natl Acad. Sci. USA, 88, 11460–11464, 1991. 

48. Campbell, R. B., Balasubramanian, S. V., and Straubinger, R. M., Physical properties of phospholipid–cationic 

lipid interactions: Influences on domain structure, liposome size and cellular uptake, 

Biochim. Biophys. Acta, 1512, 27–39, 2001. 


stability and membrane properties of paclitaxel-containing liposomes, J. Pharm. Sci., 90, 1091–1105, 

2001. 

50. Torchilin, V. P., Polymer-coated long circulating microparticulate pharmaceuticals, J. Microencapsul., 

15, 1–19, 1998. 

51. Szoka, F. and Papahadjopoulos, D., Comparative properties and methods of preparation of lipid 

vesicles (liposomes), Annu. Rev. Biophys. Bioeng., 9, 467–508, 1980. 

52. Levchenko, T. S., Rammohan, R., Lukyanov, A. N., Whiteman, K. R., and Torchilin, V. P., Liposome 

clearance in mice: The effect of a separate and combined presence of surface charge and polymer 

coating, Int. J. Pharm., 240, 95–102, 2002. 

53. Straubinger, R. M., pH-Sensitive liposomes for delivery of macromolecules into cytoplasm of 

cultured cells, Methods Enzymol., 221, 361–376, 1993. 

54. Simberg, G., Weisman, S., Talmon, Y., and Barenholz, Y., DOTAP (and other cationic lipids): 

Chemistry, biophysics, and transfection, Crit. Rev. Ther. Drug Carrier Syst., 21, 257–317, 2004. 

55. Jain, R. K., Munn, L. L., and Fukumura, D., Dissecting tumour pathophysiology using intravital 

microscopy, Nat. Rev. Cancer, 2, 266–276, 2002. 

56. Hashizume, H. et al., Openings between defective endothelial cells explain tumor vessel leakiness, 

Am. J. Pathol., 156, 1363–1380, 2000. 

57. Fukumura, D., Yuan, F., Monsky, W. L., Chen, Y., and Jain, R. K., Effect of host microenvironment 

on the microcirculation of human adenocarcinoma, Am. J. Pathol., 151, 679–688, 1997. 

58. Felgne, P. et al., Lipofection: A highly efficient, lipid-mediated DNA-transfection procedure, Proc. 

Natl Acad. Sci. USA, 84, 7413–7417, 1987. 

59. Felgner, P. L. and Ringold, G. M., Cationic liposome-mediated transfection, Nature, 337, 387–388, 

1989. 

60. Sakurai, F. et al., Effect of DNA/liposome mixing ratio on the physicochemical characteristics, 

cellular uptake and intracellular trafficking of plasmid DNA/cationic liposome complexes and subsequent 

gene expression, J. Control. Release, 255–269, 2000. 

61. Dabbas, S., Kaushik, R., Dandamudi, S., and Campbell, R.B., Physiochemical evaluation of PEGmodified 

cationic liposomes (PCLs) and properties in vitro using human primary and immortalized 

microvascular endothelial cells. J. Pharm. Sci. Submitted. 

62. Dellian, M., Yuan, F., Trubetskoy, V. S., Torchilin, V. P., and Jain, R. K., Vascular permeability in a 

human tumor xenograft: Molecular charge dependence, Br. J. Cancer, 82, 1513–1518, 2000. 

63. McLean, J. W. et al., Organ-specific endothelial cell uptake of cationic liposome–DNA complexes in 

mice, Am. J. Physiol., 273, 387–404, 1997. 

64. Tanner, F., Carr, D. P., Nabel, G. J., and Nabel, E. G., Transfection of human endothelial cells, 

Cardiovasc. Res., 35, 522–528, 1997. 

65. Matsumura, J. S., Kim, R., Shively, V. P., MacDonald, R. C., and Pearce, W. H., Characterization of 

vascular gene transfer using a novel cationic lipid, J. Surg. Res., 85, 339–345, 1999. 

66. Nabel, G. J. et al., Direct gene transfer with DNA–liposome complexes in melanoma: Expression, 

biologic activity, and lack of toxicity in humans, Proc. Natl Acad. Sci. USA, 90, 11307–11311, 1993. 

67. Pleyer, U. et al., Efficiency and toxicity of liposome-mediated gene transfer to corneal endothelial 

cells, Exp. Eye Res., 73, 1–7, 2001. 

68. Uyechi, L. S., Gagne, L., Thurston, G., and Szoka, F. C. J., Mechanism of lipoplex gene delivery in 

mouse lung: Binding and internalization of fluorescent lipid and DNA components, Gene Ther., 8, 

828–836, 2001. 

69. Trubetskoy, V. S., Torchilin, V. P., Kennel, S., and Huang, L., Cationic liposomes enhance targeted 

delivery and expression of exogenous DNA mediated by N-terminal modified poly(L-lysine)– 

antibody conjugate in mouse lung endothelial cells, Biochim. Biophys. Acta, 1131, 311–313, 1992. 



70. Abraham, N. G. et al., Transfection of the human heme oxygenase gene into rabbit coronary microvessel 

endothelial cells: Protective effect against heme hemoglobin toxicity, Proc. Natl Acad. Sci. 

USA, 92, 6798–6802, 1995. 

71. Nakatani, K. et al., Endothelial adhesion molecules in glomerular lesions: Association with their 

severity and diversity in lupus models, Kidney Int., 65, 1290–1300, 2004. 

72. Jeschke, M. G. et al., IGF-I gene transfer in thermally injured rats, Gene Ther., 6, 1015–1020, 1999. 

73. Pickering, J. G. et al., Liposome-mediated gene transfer into human vascular smooth muscle cells, 

Circulation, 89, 13–21, 1994. 

74. Tiukinhoy, S. D. et al., Development of echogenic, plasmid-incorporated, tissue-targeted cationic 

liposomes that can be used for directed gene delivery, Invest. Radiol., 35, 732–738, 2000. 

75. Hong, Z. et al., Enhanced adeno-associated virus vector expression by adenovirus protein–cationic 

liposome complex, Chin. Med. J., 108, 332–337, 1995. 

76. Torchilin, V. P., Liposomes as targetable drug carriers, CRC Crit. Rev. Ther. Drug Carrier Syst., 2, 

65–115, 1985. 

77. Ishida, T., Harashima, H., and Kiwada, H., Liposome clearance, Biosci. Reports, 22, 197–224, 2002. 

78. Harrington, K. J., Lewanski, C. R., and Stewart, J. S., Liposomes as vehicles for targeted therapy of 

cancer. Part 2. Clinical development, Clin. Oncol., 12, 16–24, 2000. 

79. Gabizon, A. and Paphadjopoulos, D., Liposome formulations with prolonged circulation time in blood 

and enhanced uptake in tumours, Proc. Natl Acad. Sci., 85, 6949–6953, 1985. 

80. Woodle, M. C. and Lasic, D. L., Sterically stabilized liposomes, Biochim. Biophys. Acta, 1113, 

171–199, 1992. 

81. Meyer, O. et al., Cationic liposomes coated with polyethylene glycol as carriers for oligonucleotides, 

J. Biol. Chem., 273, 15621–15627, 1998. 

82. Gursel, I., Gursel, M., Ishii, K. J., and Klinman, D. M., Sterically stabilized cationic liposomes 

improved the uptake and immunostimulatory activity of CpG oligonucleotides, J. Immunol., 167, 

3324–3328, 2001. 

83. Harigai, T. et al., Preferential binding of polyethylene glycol-coated liposomes containing a novel 

cationic lipid, TRX-20, to human subendothelial cells via chondroitin sulfate, Pharm. Res., 18, 

1284–1290, 2001. 

84. Strawn, L. M. and Shawver, L. K., Tyrosine kinase in disease: Overview of kinase inhibitors 

as therapeutic agents and current drugs in clinical trials, Expert Opin. Investig. Drugs, 7, 553–573, 

1998. 

85. Mendel, D. B. et al., The angiogenesis inhibitor SU5416 has long-lasting effects on vascular 

endothelial growth factor receptor phosphorylation and function, Clin. Cancer Res., 6, 4848–4858, 

2000. 

86. Brower, V., Tumor angiogenesis-new drugs on the block, Nat. Biotechnol., 17, 963–968, 1999. 

87. Hirabayashi, K., Yano, J., Takesue, H., Fujiwara, N., and Irimura, T., Inhibition of metastatic carcinoma 

cell growth in livers by poly(I):poly(C)/cationic liposome complex (LIC), Oncol. Res., 11, 

497–504, 1999. 

88. Lee, C. M. et al., Novel chondroitin sulfate-binding cationic liposomes loaded with cisplatin efficiently 

suppress the local growth and liver metastasis of tumor cells in vivo, Cancer Res., 62, 

4282–4288, 2002. 

89. Hood, J. D. and Cheresh, D. A., Role of integrins in cell invasion and migration, Nat. Rev. Cancer,2, 

91–100, 2002. 

90. Akalu, A., Cretu, A., and Brooks, P. C., Targeting integrins for control of tumour angiogenesis, Expert 

Opin. Investig. Drugs, 14, 1475–1486, 2005. 


2404–2407, 2002. 

92. Shinkarauk, S., Bayle, M., Lain, G., and Deleris, G., Vascular endothelial cell growth factor (VEGF), 

an emerging target for cancer therapy, Curr. Med. Chem. Anti-Cancer. Agents, 3, 95–117, 2003. 

93. Senger, D. R. et al., Vascular permeability factor (VPF, VEGF) in tumor biology, Metastasis Rev., 12, 

303–324, 1993. 

94. Dvorak, H. F., Brown, L. F., Detmar, M., and Dvorak, A. M., Vascular permeability factor/vascular 

endothelial growth factor, microvascular hyperpermeability and angiogenesis, Am. J. Pathol., 146, 

1029–1039, 1995. 


31 

CONTENTS 

Cell Penetrating Peptide (CPP)– 

Modified Liposomal 

Nanocarriers for Intracellular 

Drug and Gene Delivery 


31.1 Introduction ....................................................................................................................... 629 

31.2 Liposomes for Intracellular Drug Delivery ...................................................................... 630 

31.3 Transduction ...................................................................................................................... 631 

31.4 Intracellular Delivery of Pharmaceutical Nanocariers with CPPs ................................... 633 

31.5 Intracellular Delivery of Liposomes with CPPs ............................................................... 635 

References .................................................................................................................................... 637 


Although many important pharmacological targets for cancer therapy are located inside cells, 

membranes pose a serious obstacle to the uptake of charged hydrophilic molecules inside the 

cells. Many pharmacological proteins or peptides require to be intracellularly delivered to target 

and modulate cellular functions at the subcellular levels. In this connection, an important task is a 

proper identification of intracellular targets to be reached and affected. Cancer therapy provides 

a whole set of good examples to illustrate this importance. In a search of new anti-cancer drugs, a 

shift is gradually occurring from the semi-empirical approach based on evaluation of efficacy of 

candidate compounds against cultured cancer cells and animal tumor models to molecular 

mechanism-based drug discovery. 1 A huge body of information about cellular metabolic and 

signaling pathways essential for tumorogenesis and tumor cell development allows for the identification 

of proper targets for interference with the tumor growth. Quite a few molecular targets have 

already been identified. 1 The creation of a working draft of the human genome sequence 2,3 in 

combination with high-throughput methods of molecular biology promises continued rapid growth 

in identifying such targets. 4,5 However, not all drug targets that can be identified and validated by 

molecular biology tools are considered suitable for drug development, often because of their 

intracellular location. 1 Myc/Max dimerization, Src homology-2 domain interaction, and Ras/Raf 

association have been identified as promising interference targets for the inhibition of tumor 

development. 6,7 Although peptide inhibitors, discovered through the use of phage-display 8 or 

combinatorial peptide 9 libraries had excellent in vitro activity, they were not considered for drug 


629

630 

development because of poor pharmacokinetics and their inability to reach the molecular targets 

inside the cells. Sometimes, tumors result from malfunctions of tumor suppressors genes and the 

lack of activity of the proteins they encode. 5,10 In this case, the delivery into tumor cells of working 

copies of proteins obtained by recombinant methods would provide indispensable tools for validation 

of gene functions and potential development of protein or gene therapy-based methods of 

treatment. The use of these proteins for molecular target validation and eventual development of 

anti-cancer drugs is again hampered by low permeability of cell membranes. 

Intracellular transport of biologically active molecules with therapeutic properties is one of the 

key problems in drug delivery in general. However, the very nature of cell membranes prevents 

proteins, peptides, and nanoparticulate drug carriers from entering cells unless there is an active 

transport mechanism that is usually the case for very short peptides. 11 Various vector molecules 

promote the delivery of associated drugs and drug carriers inside the cells via receptor-mediated 

endocytosis. 12 This process involves the attachment of the vector molecule and an associated drug 

carrier to specific ligands on the target cell membranes, followed by the energy-dependent formation 

of endosomes. The problem, however, is that any molecule/particle entering cells via the 

endocytic pathway and becoming entrapped into endosome eventually ends in lysosome where 

active degradation processes take place under the action of numerous lysosomal enzymes. As a 

result, only a small fraction of unaffected substance appears in the cell cytoplasm. Even if an 

efficient cellular uptake via endocytosis is observed, the delivery of intact peptides and proteins 

is compromised by an insufficient endosomal escape and subsequent lysosomal degradation. 

Enhanced endosomal escape can be achieved through the use of, for example, lytic peptides, 13– 

15 pH-sensitive polymers, 16 or swellable dendritic polymers. 17 These agents have provided 

encouraging results in overcoming limitations of endocytosis-based cytoplasmic delivery, but 

there is still a need for further improvement or consideration of alternative delivery strategies. 

31.2 LIPOSOMES FOR INTRACELLULAR DRUG DELIVERY 


For almost two decades, liposomes have been considered promising carriers for biologically active 

substances. 18–20 They are biologically inert and completely biocompatible, and they cause practically 

no toxic or antigenic reactions. Drugs included into liposomes are protected from the 

destructive action of the external media, and liposomes are able to deliver their content inside 

cells and even inside different cell compartments. Water-soluble drugs can be captured by the inner 

water space of liposomes, whereas lipophilic compounds can be incorporated into the liposomal 

membrane. The definite drawback of liposomal preparations is their fast elimination from the blood 

and capture by the cells of the reticulo-endothelial system (RES), primarily, in the liver and spleen 

that is usually believed to be the result of rapid opsonization of the liposomes. 21 To increase 

liposome accumulation in the required areas, the use of targeted liposomes has been suggested. 22 

Liposomes with a specific affinity for an affected organ or tissue were believed to increase the 

efficacy of liposomal pharmaceutical agents. Immunoglobulins, primarily of the IgG class, and their 

fragments are the most promising and widely used targeting moieties for various drugs and drug 

carriers, including liposomes. Despite evident success in the development of antibody-to-liposome 

coupling techniques and improvements in the targeting efficacy, the majority of immunoliposomes 

still ended in the liver that was usually a consequence of insufficient time for the interaction 

between the target and targeted liposome. 

Different methods have been suggested to achieve long circulation of liposomes in vivo, 

including coating the liposome surface with inert, biocompatible polymers such as polyethylene 

glycol (PEG) that form a protective layer over the liposome surface and slows down the liposome 

recognition by opsonins and subsequent clearance. 23,24 Long-circulating liposomes are now widely 

used in biomedical in vitro and in vivo studies and even found their way into clinical practice. 25,26 

Explanations of the phenomenon involve the role of surface charge and hydrophilicity of PEG-coated 


CPP–Modified Liposomal Nanocarriers for Intracellular Drug and Gene Delivery 631 

liposomes, 27 the participation of PEG in the repulsive interactions between PEG-grafted membranes 

and other particles, 28 and more generally, the decreased rate of plasma protein adsorption on 

the hydrophilic surface of PEGylated liposomes. 29 Long-circulating immunoliposomes have also 

been prepared containing on their surface both antibody and PEG, thereby possessing both abilities— 

i.e., to recognize and bind the target and to circulate long enough to provide high target 

accumulation. 30 

As previously discussed, most liposomes are internalized by cells via endocytosis and destined 

to lysosomes for degradation. 31 When one needs to achieve liposome-mediated drug delivery inside 

cells into a cytoplasmic compartment, pH-sensitive liposomes are frequently used (for one of many 

reviews, see V. P. Torchilin, F. Zhou, and L. Huang 1993). 32 As noticed in Puyal et al., 33 cellular 

drug delivery mediated by pH-sensitive liposomes is not a simple intracellular leakage from the 

lipid vesicle because the drug has to also cross the endosomal membrane. It is usually supposed that 

inside an endosome, the low pH and some other factors destabilize liposomal membrane that, in 

turn, interacts with the endosomal membrane, provoking its secondary destabilization and drug 

release into the cytoplasm. To prepare pH-sensitive immunoliposomes, the latter were additionally 

supplied with surface-immobilized antibodies. The advantages of antibody and pH-sensitive liposome 

combination are mutually additive: cytoplasmic delivery, targetability, and facilitated uptake 

(i.e., improved intracellular availability) via receptor-mediated endocytosis. Successful application 

of pH-sensitive immunoliposomes has been demonstrated in the delivery of a variety of molecules 

including fluorescent dyes, anti-tumor drugs, proteins, and DNA. 34 

31.3 TRANSDUCTION 

A promising approach that seems to be the solution of overcoming the cellular barrier has emerged 

over the last decade. In this approach, certain proteins or peptides can be tethered to the hydrophilic 

drug of interest, and together, the construct possesses the ability to translocate across the plasma 

membrane and intracellularly deliver the payload. This process of translocation is called protein 

transduction. Such proteins or peptides contain domains of less than 20 amino acids, termed as 

Protein Transduction Domains (PTDs) or CPPs, that are highly rich in basic residues. These 

peptides have been used for intracellular delivery of various cargoes with molecular weights 

several times greater than their own. 35 CPPs are a group of peptides, usually containing a cluster 

of basic residues that have been recognized as promising drug delivery vectors over the last decade. 

The CPPs are gaining increased attention as they possess the remarkable property of translocating 

across the hydrophobic cell membrane that forms a formidable barrier to the entry of hydrophilic 

and high molecular-weight drugs. As a result, different therapeutic moieties that are mostly hydrophilic 

in nature and/or are of high molecular weight can be tagged to CPPs and transported across 

the cell membrane to exert their pharmacological actions at the subcellular level. This process of 

traversing across the biological barrier is called protein transduction, and it is confined to a domain 

of less than 20 amino acids, termed as PTD or CPP. CPPs are rich in basic residues and are either 

derived from corresponding transducing proteins or synthesized. The original concept of 

protein transduction came out from the observation that the trans-activating transcriptional 

activator, TAT, protein encoded by HIV-1, was efficiently internalized by cells in vitro, resulting 

in trans-activation of the viral promoter. 36,37 Subsequently, several other proteins and peptides were 

found to display the translocation activity that encompasses penetratin, 38 VP22, 39 transportan, 40 

model amphipathic peptide MAP, 41 signal sequence-based peptides, 42 and synthetic arginineenriched 

sequences. 43 

Penetratin is a CPPs derived from the homeodomain of Antennapedia (Drosophila homeoprotein). 

Homeoproteins are transcription factors that comprise a stretch of 60 amino acids called the 

homeodomain that is involved in the DNA binding. The homeodomain was shown to traverse 

through the mammalian nerve cells and accumulate in their nuclei. 44 More specifically, the 


632 


translocation ability was narrowed down to a 16-mer peptide, termed as penetratin (Antp PTD, 43– 

58 residues, Arg-Gln-Ile-Lys-Ile-Trp-Phe-Gln-Asn-Arg-Arg-Met-Lys-Trp-Lys-Lys) present in the 

third helix of the homeodomain. 38 VP22 is a herpes virus type 1 protein associated with the 

transport between cells where it also ends up in the nuclei. 39 Transportan is a chimeric CPP 

built of galanin and mastoparan. Inside the cells, transportan is taken up into the nuclei and 

concentrates in subnuclear structures, probably the nucleoli. 40 MAP is a synthetic 18-mer 

peptide capable of transporting various cargoes across the cellular membrane. MAP has presented 

the highest uptake and cargo delivery efficiency among other CPPs. 45 Signal-sequence-based 

peptides comprise membrane translocating sequences (MTSs) that direct the pre-protein toward 

the accurate intracellular organelles. Such sequences, when coupled to nuclear localization signals, 

can be directed to accumulate in the nuclei of the cells. 46 To identify the specific regions of TAT 

peptide responsible for translocation, synthetic peptides rich in arginines were prepared such as 

arginine-substituted TAT (R9-Tat) and D-amino acid substituted TAT (D-TAT). 43 Such peptides 

were internalized with similar efficiencies as TAT peptide. In addition, various arginine-rich 

peptides such as RNA-binding peptides derived from proteins HIV-1 Rev, flock house virus 

coat, and DNA-binding peptides from c-Fos, c-Jun, and yeast GCN4 also displayed 

translocation properties. 

TAT peptide (TATp) remains the most frequently used CPP for drug delivery purposes. TAT is 

transcriptional activator protein encoded by human immunodeficiency virus type 1 (HIV-1). 47 

TAT-mediated transduction was first utilized in 1994 for the intracellular delivery of variety of 

cargos such as b-galactosidase, horseradish peroxidase, RNase A, and domain III of Pseudomonas 

exotoxin A in vitro. 48 In vivo transduction using TAT peptide (37–72)-conjugated to b-galactosidase 

resulted in protein delivery to different tissues such as the heart, liver, spleen, lung, and 

skeletal muscle. Next, attempts were made to narrow down to the specific domain responsible for 

transduction. For this, synthetic peptides with deletions in the a-helix domain and the basic cluster 

domain were prepared for investigating their translocation ability. 49 The whole basic cluster from 

48 to 60 residues was found accountable for membrane translocation because any deletions or 

substitutions of basic residues in TAT (48–60) reduced the cellular uptake property. Rothbard et 

al. 50 studied TAT (48–57) by deletion analysis. They found that deletion of Gly-48 did not affect the 

transduction efficiency, whereas deletions of Lys-50,51, Arg-55–57 and Gln-54 markedly reduced 

transduction efficiencies. Therefore, the minimal transduction domain was assigned to TAT (49– 

57) residues. The commonly studied transduction domain of TAT (TAT PTD) extends from 

residues 47–57: Tyr-Gly-Arg-Lys-Lys-Arg-Arg-Gln-Arg-Arg-Arg that contain six arginines 

(Arg) and two lysine residues. 51 

Two types of the endocytic uptake of the CPPs have been proposed: the classical clathrinmediated 

endocytosis and the lipid-raft-mediated caveolae endocytosis. Clathrin-mediated endocytosis 

involves the formation of clathrin-coated membrane pits that pinch off the membrane to 

form vesicles for subsequent processing. 52 This type of endocytosis was suggested in Console et al. 

where the TAT peptide showed the co-localization with the classical endocytic marker, transferrin. 

This was substantiated further in Lundberg, Wilkstrom, and Johansson where TAT PTD and Antp 

PTD showed uptake only at 378C. In addition, the internalization required the expression of 

negatively charged glycosaminoglycans on the cell surface for interaction with CPPs that was 

followed by endocytosis. Studies also suggested that the PTDs do not provoke a real translocation, 

but they are only responsible for cell surface adherence that subsequently results in their endocytosis 

and accumulation in endosomes. 53–57 In fact, direct electrostatic interaction between the 

positive residues of CPPs and the negative residues of the cell-surface proteoglycans or glycosaminoglycans 

(such as heparan sulfate, heparin) is required in internalization regardless of the 

mechanism of cellular uptake 54,58–60 although some studies suggested the lack of correlation 

between proteoglycans and transduction process. 61 

Another proposed mechanism for internalization is caveolae-mediated clathrin-independent 

uptake. Caveolae uptake involves the formation of flask-like uncoated invaginations (50–70 nm), 



principally composed of a subclass of detergent-resistant membrane domains enriched in cholesterol 

and sphingolipids called as lipid rafts. 62 This type of uptake was suggested in. 63,64 Unlike the 

rapid uptake of transferrin, a marker for clathrin-mediated endocytosis, the internalization of TATcargo 

was very slow, reaching the plateau after several hours with the co-localization of TATp with 

the markers of caveolar uptake. The cellular uptake was affected in the presence of drugs that either 

disrupt lipid rafts or alter caveolar trafficking. 

A different mechanism has been proposed for the transport of guanidinium-rich CPPs conjugated 

to small molecules (MW!3,000 Da). 65,66 The guanidinium groups of the CPPs form 

bidentate hydrogen bonds with the negative residues on the cell-surface; the resulting ion pairs 

then translocate across the cell membrane under the influence of the membrane potential. The ion 

pair dissociates on the inner side of the membrane, releasing the CPPs into the cytosol. The number 

of guanidinium groups is critical for translocation with around eight groups being the optimum 

number for the efficient translocation. 

A recent mechanism proposed for CPP-conjugated to large cargos (MWO30,000 Da) is 

nonclathrin, noncaveolar endocytosis, called macropinocytosis. Macropinocytosis is a nonspecific 

form of cellular uptake, brought about by large vesicles known as macropinosomes that are generated 

from the actin filaments. 67 Because a majority of fusion proteins remained entrapped in 

macropinosomes, the TAT transduction domain was conjugated to a fusogenic peptide, the 

N-terminus domain of the influenza virus hemagglutinin protein HA2, to trigger the release of 

the TAT-fusion protein from endosome, enhancing their nuclear transport. A very recent study also 

demonstrates the macropinocytosis mechanism for small PTD peptides (1,000–5,000 Da). 68 

Therefore, it looks like that more than one mechanism works for CPP-mediated intracellular 

delivery of small and large molecules. Individual CPPs or CPP conjugated to small molecules are 

internalized into cells via electrostatic interactions and hydrogen bonding, whereas CPP conjugated 

to large molecules occur via the energy-dependent macropinocytosis. However, in both cases, the 

direct contact between the CPPs and negative residues on cell-surface is a requisite for 

successful transduction. 

Because the majority of the studies believe in endocytosis as a major mechanism for transduction, 

an important moment is the escape of CPPs from the endosomes and their translocation to the 

nuclei. Studies suggest that endosomal acidification prior to the disruption is required for CPP 

escape. 69 Another study suggested that the TAT-fusion proteins enter cells via the endosomal 

pathway, circumvent lysosomal degradation, and then sequester in the periphery of the nuclei. 70 

Overall, the efficiency of nuclear translocation process is limited. 

31.4 INTRACELLULAR DELIVERY OF PHARMACEUTICAL 

NANOCARIERS WITH CPPs 

The applications of TAT-mediated delivery extend to nanoparticles delivery inside cells. The 

concept of TAT-mediated nanoparticle delivery was first realized in 1999 when dextran-coated 

superparamagnetic iron oxide nanoparticles (size ca. 40 nm) conjugated to TATp (48–57), showing 

significantly higher uptake in lymphocytes in vitro than control TAT-free particle. 71 The technique 

showed potential for magnetically labeling cells in order to allow for their magnetic separation or 

magnetic resonance (MR) imaging. In vivo, TATp effectively transduced iron nanoparticles into 

hematopoietic and neural progenitor cells for stem cell analysis 72 andintoT-cellsforMR 

imaging. 73 Contrary to nuclear translocation of TATp, particulate TATp-iron conjugates were 

observed in the cytoplasm and not in the nuclei. 74 Because superparamagnetic nanoparticles 

were shown to be useful MR contrast agents for imaging and for cell labeling and cell tracking, 

the conjugation of such nanoparticles with TATp provides better signal from the treated cells for 

MR imaging. 75 For in vivo MRI, cells need to be labeled with magnetic particles through internalizing 

receptors. The limitation, however, is that most of the cells lack efficient internalization 


634 


receptors or pathways. This problem was overcome by attaching the CPP such as TAT peptide (48– 

57) to the dextran-coated superparamagnetic iron oxide particles (CLIO). The average size of the 

particles was 41 nm, and the conjugate carried average 6.7 TAT moieties per particle. The cellular 

uptake studies of TAT-CLIO were performed on mouse lymphocytes, human natural killer cells, 

and HeLa cells. In all the three cell lines tested, the uptake of TAT–CLIO nanoparticles was about 

100-fold higher than the nonmodified iron oxide particle. The labeling with TAT–CLIO did not 

induce toxicity and did not alter the differentiation or proliferation pattern of CD34C cells. The 

TAT–CLIO-labeled CD34C cells and control cells were subsequently intravenously injected into 

the immunodeficient mice. 75 Around 4% of the injected dose of the cells migrated to the bone 

marrow (per gram of the tissue), and the labeled and control cells showed similar biodistribution 

profile. Nonetheless, it was possible to visualize the labeled cells by MRI within mouse bone 

marrow at the single-cell level. Also, such magnetically labeled cells could be recovered from 

the bone marrow after in vivo homing using magnetic separation columns. 

Similarly, gold nanoparticles were also modified with TATp. 76 As with iron-conjugates, these 

particles did not reach the cell nuclei in experiments with NIH3T3 and HepG2 cells. The uptake of 

the particles was found to proceed by endocytosis. TATp (48–57) was also conjugated to FITCdoped 

silica nanoparticles (FSNPs) for bioimaging purposes. 77 TATp-modified nanoparticles have 

also been investigated for their capability to deliver the diagnostic and therapeutic agents across the 

blood–brain barrier. 78 TATp–FSNPs were prepared by microemulsion system and studied for 

labeling of human lung adenocarcinoma cells (A-549) in vitro. The cells were efficiently labeled 

with TAT–FSNPs, unlike FSNPs that showed no effective labeling. For in vivo bioimaging 

potential, TAT–FSNPs were intraarterially administered to the rats’ brains. TAT-conjugated 

FSNPs labeled the brain blood vessels, showing the potential for delivering agents to the brain 

without compromising the blood–brain barrier. 

A new application of CPP-mediated delivery is the labeling of cells with quantum dots using 

CPP-modified quantum dot-loaded polymeric micelles. 79 Quantum dots are gaining popularity over 

standard fluorophores for studying tumor pathophysiology because they are photostable and are 

very bright fluorophores. They can be tuned to a narrow emission spectrum, and they are relatively 

insensitive to the wavelength of the excitation light. Quantum dots have the ability to distinguish 

tumor vessels from both the perivascular cells and the matrix with concurrent imaging. Quantum 

dots were trapped within micelles prepared of PEG–phosphatidyl ethanolamine (PEG–PE) conjugates 

bearing TAT–PEG–PE linker. TAT-quantum dot conjugates could label mouse endothelial 

cells in vitro. For in vivo racking, bone marrow-derived progenitor cells were labeled with TATbearing 

quantum dot-containing micelle ex vivo and then injected in the mouse bearing tumor in a 

cranial window model. It was then possible to track the movement of labeled progenitor cell to 

tumor endothelium, introducing an attempt toward understanding fine details of tumor 

neovascularization. 

CPPs have also been used to enhance the delivery of genes via solid lipid nanoparticle (SLN). 80 

SLN gene vector was modified with dimeric TATp (TAT2) and compared with polyethylenimine 

(PEI) for gene expression in vitro and in vivo. The presence of TAT2 in SLN gene vector enhanced 

the gene transfection compared to PEI both in vitro and in vivo. In another study, TATp (47–57) 

conjugated to nanocage structures showed binding and transduction of the cells in vitro. 81 The shell 

cross-linked (SCK) nanoparticles were prepared by the micellization of amphiphilic block copolymers 

of poly(epsilon-caprolactone-b-acrylic acid) and conjugated to TATp that was 

independently built on a solid support, resulting in TAT-modified nanocage conjugate. Such 

conjugate was analyzed by confocal microscopy with CHO and HeLa cells. The conjugated nanoparticles 

showed binding and transduction inside the cells. The authors then characterized the SCK 

nanoparticles for the optimum number of TATp per particle required to enhance the transduction 

efficiency. 82 The authors prepared the conjugates with 52, 104, and 210 CPP peptides per particle 

that were then evaluated for the biocompatibility in vitro and in vivo. 83 In vitro studies showed the 

inflammatory responses to the conjugates, but in vivo evaluation of the conjugates in mice did not 



result in major incompatible responses. Therefore, TATp–SCK conjugate can be used as scaffolds 

for preparing antigen for immunization. 

TAT-mediated nanoparticulate delivery is also finding its way in vaccination fields to elicit 

better immune response. When TAT (1–72)-coated anionic nanoparticles were used to immunize 

mice, it generated antibodies and T helper type-1 immune response to TAT. 84 In another study, 

TAT microspheres were used for vaccinations. 85 Anionic microspheres of different compositions, 

size, and surface charge density were prepared, and all of them adsorbed biologically active TAT 

protein in a reversible mode. The microspheres were intracellularly delivered by TAT and were not 

toxic, both in vitro and in vivo. 

31.5 INTRACELLULAR DELIVERY OF LIPOSOMES WITH CPPs 

CPPs have also augmented the delivery of liposomal drug carriers. TAT peptide (47–57)-modified 

liposomes could be intracellularly delivered in different cells such as murine Lewis lung carcinoma 

(LLC) cells, human breast tumor BT20 cells, and rat cardiac myocyte H9C2 cells. 86 The liposomes 

were tagged with TAT peptide via the spacer p-nitrophenylcarbonyl–PEG–PE at the density of 500 

TAT peptide per single liposome vesicle. It was shown that the cells treated with liposomes where 

TAT peptide–cell interaction was hindered either by direct attachment of TAT peptide to the 

liposome surface or by the long PEG grafts on the liposome surface shielding the TAT moiety 

did not show TAT-liposome internalization; however, the preparations of TAT-liposomes that 

allowed for the direct contact of TAT peptide residues with cells displayed an enhanced uptake 

by the cells. This suggested that the translocation of TAT peptide (TATp)-liposomes into cells 

requires direct free interaction of TAT peptide with the cell surface. Further studies on the intracellular 

trafficking of rhodamine-labeled TATp-liposomes loaded with FITC-dextran revealed that 

TATp-liposomes remained intact inside the cell cytoplasm within 1 h of translocation; after 2 h, 

they migrated into the perinuclear zone, and eventually, the liposomes disintegrated there. 87 

The TATp-liposomes were also investigated for their gene delivery ability. For this, TATpliposomes 

prepared with the addition of a small quantity of a cationic lipid (DOTAP) were 

incubated with DNA. The liposomes formed firm noncovalent complexes with DNA. Such 

TATp-liposome-DNA complexes, when incubated with mouse fibroblast NIH 3T3 and cardiac 

myocytes H9C2, showed substantially higher transfection in vitro with lower cytotoxicity than 

the commonly used Lipofectin w . NIH/3T3, BT20, or H9C2 cells were transfected by incubation 

with TAT peptide–liposome/plasmid complexes in serum-free media for 4 h at 378C under 6% 

CO 2. The flow cytometry data demonstrated that the treatment of NIH/3T3 cells with TAT peptide– 

liposome/pEGFP-N1 complexes results in high fluorescence, i.e., high transfection outcome. 

Similar results were obtained with all cell lines tested. Confocal microscopy confirmed the transfection 

of both HCC150 and H9C2 cells with TAT peptide–liposome/DNA complexes 

(Figure 31.1). From 30 to 50% of both cell types in the field of view show a bright green fluorescence, 

whereas lower fluorescence was observed in virtually all cells. Under in vivo conditions, 

the intratumoral injection of TATp-liposome–DNA complexes into the LLC tumor in mice resulted 

in an efficient transfection of the tumor cells. Histologically, hematoxylin/eosin-stained tumor 

slices in both control and experimental animals showed a typical pattern of poorly differentiated 

carcinoma; however, under the fluorescence microscope, samples from control mice (nontreated 

mice or mice injected with TAT peptide–free liposome/plasmid complexes) showed only a background 

fluorescence, and slices from tumors injected with TAT peptide–liposome/plasmid 

complexes contained bright green fluorescence in tumor cells, indicating an efficient TAT 

peptide-mediated transfection. The study implicated the usefulness of TATp-liposomes for 

in vitro and localized in vivo gene therapy. 

Another study examined the kinetics of uptake of the TAT- and penetratin-modified liposomes. 

88 It was found that the translocation of liposomes by TAT peptide or penetratin was 


636 

FIGURE 31.1 The enhancement of the liposome-mediated transfection by TAT peptide. HCC1500 cells 

(human breast carcinoma; a,b) or H9C2 cells (murine myoblasts; c) were incubated with the presence of 

liposome/pEGFP-N1 plasmid (encoding for the Green Fluorescence Protein, GFP) complexes (10 mg DNA per 

100,000 cells) for 4 h at 378C. The transfection efficiency (the appearance of the green fluorescence of the GFP 

inside cells) was detected after 72 h. (a), Cells were incubated with control plain (TAT-free) plasmid-bearing 

liposomes; (b,c), Cells were incubated with plasmid-bearing TAT-liposomes of the same composition. (a), (b) 

and (c)—fluorescent microscopy with FITC filter. Background transfection can only be seen with controls, 

whereas the introduction of TAT peptide into the preparation provides a dramatic enhancement of the GFP 

expression, i.e., increases the transfection efficiency. 

proportional to the number of peptide molecules attached to the liposomal surface. A peptide 

number of as few as five was already sufficient to enhance the intracellular delivery of liposomes. 

The kinetics of the uptake were peptide- and cell-type dependent. With TATp-liposomes, the 

intracellular accumulation was time-dependent, and with penetratin-liposomes, the accumulation 

within the cells was quick to reach the peak within 1 h, after that, it gradually declined. 

A study on the similar lines showed that Antp (43–58) and TAT (47–57) peptides coupled to 

small unilamellar liposomes were accumulated in higher proportions within tumor cells and 

dendritic cells than unmodified control liposomes. 89 The uptake was time- and concentrationdependent, 

and at least, 100 PTD molecules per small unilamellar liposomes were required for 

efficient translocation inside cells. The uptake of the modified liposomes was inhibited by the 

preincubation of liposomes with heparin, confirming the role of heparan sulfate proteoglycans in 

CPP-mediated uptake. 

In a different approach for improving the transfection and protecting DNA from degradation, 

thiocholesterol-based cationic lipids (TCL) were used in the formation of nanolipoparticles (NLPs). 

The NLPs were sequentially modified with TAT peptide that resulted in TAT–NLPs with a zwitterionic 

surface and higher transfection efficiency than for the cationic NLPs. 90 

TABLE 31.1 

Intracellular Delivery of Nanoparticles by CPPs 

Particle and size CPP Cell 

CLIO (MION) particles, 

40 nm 


TATp Mouse lymphocytes, human natural killer, HeLa, human hematopoietic 

CD34C, mouse neural progenitor C17.2, human lymphocytes CD4C, 

T-cells, B-cells, macrophages 

Gold particles, 20 nm TATp NIH3T3, HepG2, HeLa, human fibroblast HTERT-BJ1 

Quantum dot-loaded 

polymeric micelles, 

20 nm 

TATp Mouse endothelial cells, bone marrow-derived progenitor cells 

Sterically stabilized 

liposomes, 200 nm 

TATp Mouse LLC, human BT20, rat H9C2, LLC tumor in mice 

Sterically stabilized TATp or Human bladder carcinoma HTB-9, murine colon carcinoma C26, human 

liposomes, 65–75 nm penetratin epidermoid carcinoma A431, human breast cancer SK-BR-3, MCF7/WT, 

MCF7/ADR, murine bladder cancer MBT2, dendritic cells 



Although initial studies suggested an energy-independent character for the internalization of 

TATp-liposomes, 86 recent studies have revealed the endocytosis as the main mechanism for the 

intracellular uptake of TATp-liposomes. As was shown in M. M. Fretz, et al., 93 the conjugation of 

TAT peptide to lipoplexes enhanced the gene transfection in primary cell cultures by the endocytic 

uptake. Similarly, coupling of TAT peptide to the outer surface of liposomes resulted in an 

enhanced binding and endocytosis of the liposomes in ovarian carcinoma cells. 92 The binding 

was inhibited in the presence of heparin or dextran sulfate, suggesting that the proteoglycans 

expressed on the cell surface are also involved in cell binding in this case. In contrast, a new 

class of transducing peptides, haptides, after binding to the liposomal surface, augmented liposomes 

penetration through the cell membrane into the cell cytoplasm by a nonreceptor mediated 

process, 93 confirming that a variety of mechanisms could be involved in CPP-mediated intracellular 

delivery of nanoparticulates. 

CPPs clearly can serve as versatile delivery vectors for intracellular drug delivery. They can 

bring inside cells a wide range of cargos of different sizes—from small molecules to relatively large 

nanoparticles (see Table 31.1 for nanoparticles). Of the different CPPs used, TATP remains the 

most frequently used for this purpose in experimental cancer therapy. Intracellular cytoplasmic 

delivery of drugs and DNA by CPP-modified pharmaceutical nanocarriers may find applications for 

in vitro and ex vivo cell treatment as well as in different protocols of local drug application. 

REFERENCES 

1. Gibbs, J. B., Mechanism-based target identification and drug discovery in cancer research, Science, 

287, 1969–1973, 2000. 

2. Venter, J. C., Adams, M. D., Myers, E. W., Li, P. W., Mural, R. J., Sutton, G. G., Smith, H. O., 

Yandell, M., Evans, C. A., and Holt, R. A., The sequence of the human genome, Science, 291, 

1304–1351, 2001. 

3. Lander, E. S., Linton, L. M., Birren, B., Nusbaum, C., Zoby, M. C., Baldwin, M.C, Devon, K., Dewar, 

K., Doyle, M., and Fitzhugh, W., Initial sequencing and analysis of the human genome, Nature, 409, 

860–921, 2001. 

4. Workman, P., New drug targets for genomic cancer therapy: Successes, limitations, opportunities and 

future challenges, Curr. Cancer Drug Targets, 1, 33–47, 2001. 

5. Balmain, A., Cancer genetics: From Boveri and Mendel to microarrays, Nat. Rev. Cancer, 1, 77–82, 

2001. 

6. Gibbs, J. B. and Oliff, A., Pharmaceutical research in molecular oncology, Cell, 79, 193–198, 1994. 

7. Sawyer, T. K., Src homology-2 domains: Structure, mechanisms, and drug discovery, Biopolymers, 

47, 243–261, 1998. 

8. Katz, B. A., Structural and mechanistic determinants of affinity and specificity of ligands discovered 

or engineered by phage display, Annu. Rev. Biophys. Biomol. Struct., 26, 27–45, 1997. 

9. Cortese, R., Ed, Combinatorial Libraries: Synthesis, Screening and Application Potential, Walter de 

Gruyter, Inc., New York, 1995. 

10. Hussain, S. P., Hofseth, L. J., and Harris, C. C., Tumor suppressor genes: At the crossroads of 

molecular carcinogenesis, molecular epidemiology and human risk assessment, Lung Cancer, 34 

(Suppl. 2), S7–S15, 2001. 

11. Egleton, R. D. and Davis, T. P., Bioavailability and transport of peptides and peptide drugs into the 

brain, Peptides, 18, 1431–1439, 1997. 

12. Park, J. W., Kirpotin, D. B., Hong, K., Shalaby, R., Shao, Y., Nielsen, U. B., and Marks, J. D., Tumor 

targeting using anti-her2 immunoliposomes, J. Controlled Release, 74, 95–113, 2001. 

13. Kamata, H., Yagisawa, H., Takahashi, S., and Hirata, H., Amphiphilic peptides enhance the efficiency 

of liposome-mediated DNA transfection, Nucleic Acids Res., 22, 536–537, 1994. 

14. Midoux, P., Kichler, A., Boutin, V., Maurizot, J. C., and Monsigny, M., Membrane permeabilization 

and efficient gene transfer by a peptide containing several histidines, Bioconjugate Chem., 9, 260–267, 

1998. 


638 


15. Mastrobattista, E., Koning, G. A., Van Bloois, L., Filipe, A. C. S., Jiskoot, W., and Storm, G., 

Functional characterization of an endosome-disruptive peptide and its application in cytosolic 

delivery of immunoliposome-entrapped proteins, J. Biol. Chem., 277, 27135–27143, 2002. 

16. Lackey, C. A., Press, O., Hoffman, A., and Styton, P., A biomimetic pH-responsive polymer directs 

endosomal release and intracellular delivery of an endocytosed antibody complex, Bioconjugate 

Chem., 13, 996–1001, 2002. 

17. Padilla De Jesus, O. L., Ihre, H. R., Gagne, L., Frechet, J. M., and Szoka, F. C. Jr., Polyester dendritic 

systems for drug delivery applications: In vitro and in vivo evaluation, Bioconjugate Chem., 13, 

453–461, 2002. 

18. Gregoriadis, G., Ed, Liposomes as Drug Carriers, Wiley, Chinchester, UK, 1988. 

19. Lasic, D. D., Liposomes. From Physics to Applications, Elsevier, Amsterdam, 1993. 

20. Lasic, D. D. and Papahadjopoulos, D., Eds., Medical Application of Liposomes, Elsevier, Amsterdam, 

1998. 

21. Senior, J. H., Fate and behavior of liposomes in vivo: A review of controlling factors, Crit. Rev. Ther. 

Drug Carrier Syst., 3, 123–193, 1987. 

22. Torchilin, V. P., Liposomes as targetable drug carriers, CRC Crit. Rev. Ther. Drug Carrier Syst., 1, 

65–115, 1985. 

23. Klibanov, A. L., Maruyama, K., Torchilin, V. P., and Huang, L., Amphipatic polyethyleneglycols 

effectively prolong the circulation time of liposomes, FEBS Lett., 268, 235–238, 1990. 

24. Torchilin, V. P. and Trubetskoy, V. S., Which polymers can make nanoparticulate drug carriers longcirculating?, 


25. Martin, F. and Lasic, D., Eds., Stealth w Liposomes, CRC Press, Boca Raton, 1995. 


chemotherapy, Cancer Invest., 19, 424–436, 2001. 

27. Blume, G. and Cevc, G., Molecular mechanism of the lipid vesicle longevity in vivo, Biochim. 

Biophys. Acta, 1146, 157–168, 1993. 

28. Kenworthy, A. K., Hristova, K., Needham, D., and McIntosh, T. J., Range and magnitude of the steric 

pressure between bilayers containing phospholipids with covalently attached poly(ethylene glycol), 

Biophys. J., 68, 1921–1936, 1995. 

29. Woodle, M. C., Newman, M. S., and Cohen, J. A., Sterically stabilized liposomes: Physical and 

biological properties, J. Drug Target., 5, 397–403, 1994. 

30. Torchilin, V. P., Klibanov, A. L., Huang, L., O’Donnell, S., Nossiff, N. D., and Khaw, B. A., Targeted 

accumulation of polyethelene glycol-coated immunoliposomes in infarcted rabbit myocardium, 

FASEB J., 6, 2716–2719, 1992. 

31. Straubinger, R. M. N., pH-sensitive liposomes mediate cytoplasmic delivery of encapsulated macromolecules, 

FEBS Lett., 179, 148–154, 1985. 

32. Torchilin, V. P., Zhou, F., and Huang, L., pH-Sensitive liposomes (review), J. Liposome Res., 3, 

201–255, 1993. 

33. Puyal, C., Milhaud, P., Bienvenue, A., and Philippot, J. R., A new cationic liposome encapsulating 

genetic material. A potential delivery system for polynucleotides, Eur. J. Biochem., 228, 697–703, 

1995. 

34. Collins, D. and Huang, L., Cytotoxicity of diphtheria toxin A fragment to toxin-resistant murine cells 

delivered by pH-sensitive immunoliposomes, Cancer Res., 47, 735–739, 1987. 

35. Zorko, M. and Langel, U., Cell-penetrating peptides: Mechanism and kinetics of cargo delivery, Adv. 

Drug Deliv. Rev., 57, 529–545, 2005. 

36. Green, M. and Loewenstein, P. M., Autonomous functional domains of chemically synthesized 

human immunodeficiency virus tat trans-activator protein, Cell, 55, 1179–1188, 1988. 

37. Frankel, A. D. and Pabo, C. O., Cellular uptake of the tat protein from human immunodeficiency virus, 

Cell, 55, 1189–1193, 1988. 

38. Derossi, D., Joliot, A. H., Chassaing, G., Derossi, D., and Prochiantz, A., The third helix of the 

Antennapedia homeodomain translocates through biological membranes, J. Biol. Chem., 269, 

10444–10450, 1994. 

39. Elliot, G. and O’Harre, P., Intercellular trafficking and protein delivery by a herpesvirus structural 

protein, Cell, 88, 223–233, 1997. 



40. Pooga, M., Hällbrink, M., Zorko, M., and Langel, Ü, Cell penetration by transportan, FASEB J., 12, 

67–77, 1998. 

41. Oehlke, J., Scheller, A., Wiesner, B., Krause, E., Beyermann, M., Klauschenz, E., Melzig, M., and 

Bienert, M., Cellular uptake of an alpha-helical amphipathic model peptide with the potential to 

deliver polar compounds into the cell interior non-endocytically, Biochim. Biophys. Acta, 1414, 

127–139, 1998. 

42. Lindgren, M., Hällbrink, M., Prochiantz, A., and Langel, U., Cell-penetrating peptides, Trends Pharmacol. 

Sci., 21, 99–103, 2000. 

43. Futaki, S., Suzuki, T., Ohashi, W., Yagami, T., Tanaka, S., Ueda, K., and Sugiura, Y., Arginine-rich 

peptides. An abundant source of membrane-permeable peptides having potential as carriers for 

intracellular protein delivery, J. Biol. Chem., 276, 5836–5840, 2001. 

44. Joliot, A., Pernelle, C., Deagostini-Bazin, H., and Prochiantz, A., Antennapedia homeobox peptide 

regulates neural morphogenesis, Proc. Natl. Acad. Sci. USA, 88, 1864–1868, 1991. 

45. Hällbrink, M., Florén, A., Elmquist, A., Pooga, M., Bartfai, T., and Langel, U., Cargo delivery kinetics 

of cell-penetrating peptides, Biochim. Biophys. Acta, 1515, 101–109, 2001. 

46. Lin, Y. Z., Yao, S. Y., Veach, R. A., Torgerson, T. R., and Hawiger, J., Inhibition of nuclear 

translocation of transcription factor NF-kappa B by a synthetic peptide containing a cell 

membrane-permeable motif and nuclear localization sequence, J. Biol. Chem, 270, 14255–14258, 

1995. 

47. Calnan, B. J., Tidor, B., Biancalana, S., Hudson, D., and Frankel, A. D., Arginine-mediated RNA 

recognition: The arginine fork, Science, 252, 1167–1171, 1991. 

48. Schwarze, S. R., Hruska, K. A., and Dowdy, S. F., Protein transduction: Unrestricted delivery into all 

cells?, Trends Cell Biol., 10, 290–295, 2000. 

49. Mitchell, D. J., Kim, D. T., Steinman, L., Fathman, C. G., and Rothbard, J. B., Polyarginine enters 

cells more efficiently than other polycationic homopolymers, J. Pept. Res., 56, 318–325, 2000. 

50. Rothbard, J. B., Kreider, E., Vandeusen, C. L., Wright, L., Wylie, B. L., and Wender, P. A., Argininerich 

molecular transporters for drug delivery: Role of backbone spacing in cellular uptake, J. Med. 

Chem., 45, 3612–3618, 2002. 

51. Sieczkarski, S. B. and Whittaker, G. R., Dissecting virus entry via endocytosis, J. Gen. Virol., 83, 

1535–1545, 2002. 

52. Richard, J. P., Melikov, K., Vives, E., Ramos, C., Verbeure, B., Gait, M. J., Chernomordik, L. V., and 

Lebleu, B., Cell-penetrating peptides. A reevaluation of the mechanism of cellular uptake, J. Biol. 

Chem., 278, 585–590, 2003. 

53. Console, S., Marty, C., Garcia-Echeverria, C., Schwendener, R., and Ballmer-Hofer, K., Antennapedia 

and HIV transactivator of transcription (TAT) protein transduction domains promote 

endocytosis of high molecular weight cargo upon binding to cell surface glycosaminoglycans, 

J. Biol. Chem., 278, 35109–35114, 2003. 

54. Lundberg, M., Wilkstrom, S., and Johansson, M., Cell surface adherence and endocytosis of protein 

transduction domains, Mol. Ther., 8, 143–150, 2003. 

55. Leifert, J. A., Harkins, S., and Whitton, J. L., Full-length proteins attached to the HIV tat protein 

transduction domain are neither transduced between cells, nor exhibit enhanced immunogenicity, 

Gene Ther., 9, 1422–14428, 2002. 

56. Vives, E., Richard, J. P., Rispal, C., and Lebleu, B., TAT peptide internalization: Seeking the 

mechanism of entry, Curr. Protein Pept. Sci., 4, 125–132, 2003. 

57. Tyagi, M., Rusnati, M., Presta, M., and Giacca, M., Internalization of HIV-1 tat requires cell surface 

heparan sulfate proteoglycans, J. Biol. Chem., 276, 3254–3261, 2001. 

58. Sandgren, S., Cheng, F., and Belting, M., Nuclear targeting of macromolecular polyanions by an HIV- 

Tat derived peptide. Role for cell-surface proteoglycans, J. Biol. Chem, 277, 38877–38883, 2002. 

59. Mai, J. C., Shen, H., Watkins, S. C., Cheng, T., and Robbins, P. D., Efficiency of protein transduction 

is cell type-dependent and is enhanced by dextran sulfate, J. Biol. Chem., 277, 30208–30218, 2002. 

60. Violini, S., Sharma, V., Prior, J. L., Dyszlewski, M., and Piwnica-Worms, D., Evidence for a plasma 

membrane-mediated permeability barrier to Tat basic domain in well-differentiated epithelial cells: 

Lack of correlation with heparan sulfate, Biochemistry, 41, 12652–12661, 2002. 

61. Anderson, R. G., The caveolae membrane system, Annu. Rev. Biochem., 67, 199–225, 1998. 


640 


62. Fittipaldi, A., Ferrari, A., Zoppe, M., Arcangeli, C., Pellegrini, V., Beltram, F., and Giacca, M., Cell 

membrane lipid rafts mediate caveolar endocytosis of HIV-1 Tat fusion proteins, J. Biol. Chem., 278, 

34141–34149, 2003. 

63. Ferrari, A., Pellegrini, V., Arcangeli, C., Fittipaldi, A., Giacca, M., and Beltram, F., Caveolaemediated 

internalization of extracellular HIV-1 tat fusion proteins visualized in real time, Mol. 

Ther., 8, 284–294, 2003. 

64. Rothbard, J. B., Jessop, T. C., and Wender, P. A., Role of membrane potential and hydrogen bonding 

in the mechanism of translocation of guanidinium-rich peptides into cells, J. Am. Chem. Soc., 126, 

9506–9507, 2004. 

65. Rothbard, J. B., Jessop, T. C., and Wender, P. A., Adaptive translocation: The role of hydrogen 

bonding and membrane potential in the uptake of guanidinium-rich transporters into cells, Adv. 

Drug Deliv. Rev., 57, 495–504, 2005. 

66. Swanson, J. A. and Watts, C., Macropinocytosis, Trends Cell Biol., 5, 424–428, 1995. 

67. Kaplan, I. M., Wadia, J. S., and Dowdy, S. F., Cationic TAT peptide transduction domain enters cells 

by macropinocytosis, J. Controlled Release, 102, 247–253, 2005. 

68. Potocky, T. B., Menon, A. K., and Gellman, S. H., Cytoplasmic and nuclear delivery of a TATderived 

peptide and a beta-peptide after endocytic uptake into HeLa cells, J. Biol. Chem., 278, 

50188–50194, 2003. 

69. Caron, N. J., Quenneville, S. P., and Tremblay, J. P., Endosome disruption enhances the functional 

nuclear delivery of Tat-fusion proteins, Biochem. Biophys. Res. Commun., 319, 12–20, 2004. 

70. Zhao, M., Kircher, M. F., Josephson, L., and Weissleder, R., Differential conjugation of tat peptide to 

superparamagnetic nanoparticles and its effect on cellular uptake, Bioconjugate Chem., 13, 840–844, 

2002. 

71. Lewin, M., Carlesso, N., and Tung, C. H., Tat peptide-derivatized magnetic nanoparticles allow 

in vivo tracking and recovery of progenitor cells, Nat. Biotechnol., 18, 410–414, 2000. 

72. Dodd, C. H., Hsu, H. C., Chu, W. J., Yang, P., Zhang, H. G., Zhang, J. D. Jr., Zinn, K., Forder, J., 

Josephson, L., Weissleder, R., Mountz, J. M., and Mountz, J. D., Normal T-cell response and in vivo 

magnetic resonance imaging of T cells loaded with HIV transactivator-peptide-derived superparamagnetic 

nanoparticles, J. Immunol. Methods, 256, 89–105, 2001. 

73. Kaufman, C. L., Williams, M., Ryle, L. M., Smith, T. L., Tanner, M., and Ho, C., Superparamagnetic 

iron oxide particles transactivator protein-fluorescein isothiocyanate particle labeling for in vivo 

magnetic resonance imaging detection of cell migration: Uptake and durability, Transplantation, 

76, 1043–1046, 2003. 

74. Zhao, M., Kircher, M. F., Josephson, L., and Weissleder, R., Differential conjugation of tat peptide to 

superparamagnetic nanoparticles and its effect on cellular uptake, Bioconjugate Chem., 13, 840–844, 

2002. 

75. Tkachenko, A. G., Xie, H., Liu, Y. L., Coleman, D., Ryan, J., Glomm, W. R., Shipton, M. K., and 

Feldheim, D. L., Cellular trajectories of peptide-modified gold particle complexes: Comparison of 

nuclear localization signals and peptide transduction domains, Bioconjugate Chem., 15, 482–490, 

2004. 

76. Santra, S., Yang, H., Dutta, D., Stanley, J., Holloway, P., Tan, W., Moudgil, B., and Mericle, R., TAT 

conjugated, FITC doped silica nanoparticles for bioimaging applications, Chem. Commun. (Camb.), 

24, 2810–2811, 2004. 

77. Santra, S., Yang, H., Stanley, J. T., Holloway, P. H., Moudgil, B. M., Walter, G., and Mericle, R. A., 

Rapid and effective labeling of brain tissue using TAT-conjugated CdS:Mn/ZnS quantum dots, Chem. 

Commun. (Camb.), 25, 3144–3146, 2005. 

78. Stroh, M., Zimmer, J. P., Zimmer, D. G., Levchenko, T. S., and Cohen, K. S., Quantum dots spectrally 

distinguish multiple species within the tumor milieu in vivo, Nat. Med., 11, 678–682, 2005. 

79. Rudolph, C., Schillinger, U., Ortiz, A., Tabatt, A., Plank, C., Müller, R. H., and Rosenecker, J., 

Application of novel solid lipid nanoparticle (SLN)-gene vector formulations based on a dimeric 

HIV-1 TAT-peptide in vitro and in vivo, Pharm. Res., 21, 1662–1669, 2004. 

80. Liu, J., Zhang, Q., Remsen, E. E., and Wooley, K. L., Nanostructured materials designed for cell 

binding and transduction, Biomacromolecules, 2, 362–368, 2001. 

81. Becker, M. L., Remsen, E. E., Pan, D., and Wooley, K. L., Peptide-derivatized shell-cross-linked 

nanoparticles. 1. Synthesis and characterization, Bioconjugate Chem., 15, 699–709, 2004. 



82. Becker, M. L., Bailey, L. O., and Wooley, K. L., Peptide-derivatized shell-cross-linked nanoparticles. 

2. Biocompatibility evaluation, Bioconjugate Chem., 15, 710–717, 2004. 

83. Cui, Z., Patel, J., Tuzova, M., Ray, P., Phillips, R., Woodward, J. G., Nath, A., and Mumper, R. J., 

Strong T cell type-1 immune responses to HIV-1 Tat (1-72) protein-coated nanoparticles, Vaccine, 22, 

2631–2640, 2004. 

84. Caputo, A., Brocca-Cofano, E., Castaldello, A., De Michele, R., Altavilla, G., Marchisio, M., Gavioli, 

R., Rolen, V., Chiarantini, L., Cerasi, A., Magnani, M., Cafaro, A., Sparnacci, K., Laus, M., Tondalli, 

L., and Ensoli, B., Novel biocompatible anionic polymeric microspheres for the delivery of the HIV-1 

Tat protein for vaccine application, Vaccine, 22, 2910–2924, 2004. 

85. Torchilin, V. P., Rammohan, R., Weissig, V., and Levchenko, T. S., TAT peptide on the surface of 

liposomes affords their efficient intracellular delivery even at low temperature and in the presence of 

metabolic inhibitors, Proc. Natl. Acad. Sci. USA, 98, 8786–8796, 2001. 

86. Torchilin, V. P., Levchenko, T. S., Rammohan, R., and Volodina, B., Cell transfection in vitro and 

in vivo with nontoxic TAT peptide-liposome-DNA complexes, Proc. Natl. Acad. Sci. USA, 100, 

1972–1977, 2003. 

87. Tseng, Y. L., Liu, J. J., and Hong, R. L., Translocation of liposomes into cancer cells by cellpenetrating 

peptides penetratin and tat: A kinetic and efficacy study, Mol. Pharmacol., 62, 

864–872, 2002. 

88. Marty, C., Meylan, C., Schott, H., Ballmer-Hofer, K., and Schwendener, R. A., Enhanced heparan 

sulfate proteoglycan-mediated uptake of cell-penetrating peptide-modified liposomes, Cell Mol. Life 

Sci., 61, 1785–1794, 2004. 

89. Huang, Z., Li, W., MacKay, J. A., and Szoka, F. C. Jr., Thiocholesterol-based lipids for ordered 

assembly of bioresponsive gene carriers, Mol. Ther., 11, 409–417, 2005. 

90. Hyndman, L., Lemoine, J. L., Huang, L., Porteous, D. J., and Boyd, A. C., HIV-1 Tat protein 

transduction domain peptide facilitates gene transfer in combination with cationic liposomes, 

J. Controlled Release, 99, 435–444, 2004. 

91. Fretz, M. M., Koning, G. A., Mastrobattista, E., Jiskoot, W., and Storm, G., OVCAR-3 cells internalize 

TAT-peptide modified liposomes by endocytosis, Biochim. Biophys. Acta, 1665, 48–56, 2004. 

92. Gorodetsky, R., Levdansky, L., Vexler, A., Shimeliovich, I., Kassis, I., Ben-Moshe, M., Magdassi, S., 

and Marx, G., Liposome transduction into cells enhanced by haptotactic peptides (Haptides) homologous 

to fibrinogen C-termini, J. Controlled Release, 95, 477–488, 2004. 


32 

CONTENTS 

RGD-Modified Liposomes 

for Tumor Targeting 

P. K. Dubey, S. Mahor, and S. P. Vyas 

32.1 Introduction ....................................................................................................................... 643 

32.2 Tumor Vasculature Targeting ........................................................................................... 644 

32.2.1 Rationale for Tumor Vasculature Targeting....................................................... 644 

32.2.2 Molecular Targets in Angiogenic Tumor Vasculature....................................... 646 

32.3 Role of RGD in Integrin- and Fibronectin-Mediated Bioevents...................................... 647 

32.3.1 Integrins ............................................................................................................... 647 

32.3.2 Fibronectin........................................................................................................... 650 

32.3.3 RGD Mediated Metastasis Inhibition ................................................................. 651 

32.4 RGD-Modified Liposomes for Cancer Therapeutics........................................................ 653 

32.5 RGD-Modified Liposomes in Cancer Gene Therapy ....................................................... 656 

References .................................................................................................................................... 657 


Various drug delivery systems have been developed or are under development in order to minimize 

drug degradation or loss, to prevent harmful side effects, and to increase drug bioavailability and the 

fraction of the drug accumulated in the required zone. Among drug carriers, one can name microparticles, 

nanoparticles, synthetic polymers, cell ghosts, lipoproteins, micelles, niosomes, and 

liposomes. Each of these carriers offers its own advantages and has its own shortcomings; therefore, 

the choice of a certain carrier for each given case can be made only taking into account the 

whole bunch of relevent considerations. These carriers can be made slowly biodegradable, 

stimuli-sensitive, e.g., pH- or temperature sensitive, and even targeted, e.g., by conjugating them 

with various ligands. 

Liposomes-encapsulated anti-cancer drugs reveal their potential for increased therapeutic efficacy 

and decreased nonspecific toxicities as a result of their ability to enhance the delivery of 

chemotherapeutic agents selectively or preferentially to tumors (Papahadjopoulos and Gabizon 

1995; Martin 1998). However, these liposomes localize in the tumor extracellular compartment 

and are rarely taken up by the cancer cells located deep within the tumor (Papahadjopoulos et al. 

1991a, 1991b; Gabizon and Papahadjopoulos 1992) so that these cancer cells are readily reached 

with high concentration of drug and are given an opportunity to opt for drug-resistance. To further 

enhance cytotoxic efficacy, selective delivery of drugs to target cells can be achieved through 

liposomes appended with antibodies (Maruyama et al. 1990; Lopes de Menezes, Pilarski, and 


643

644 

Allen 1998), serum proteins (Brown and Silvius 1990; Lundberg, Hong, and Papahadjoppoulos 

1993), or antibody fragments (Park et al. 1995; Kirpotin et al. 1997) that categorically recognize 

specific determinants (receptors) on target cells. However, attachment of antibodies and larger 

proteins to the exterior of liposomes has not been without problems. Chemistries producing phospholipid 

headgroup-protein coupling are often complex, jeopardizing the subsequent use of such 

engineered liposomes, especially in living systems. These chemistries can also change the protein’s 

confirmation and do not guarantee the final orientation with respect to either the potential ligand or 

the phospholipid bilayer, thereby affecting the effectiveness of the targeting system. 

An alternative to the coupling of larger proteins is an anchoring of a defined small peptide 

domain on to the liposome surface (Gyongyossy-Issa, Muller, and Devine 1998). Ligand targeting 

using small peptides has advantages over the use of large protein molecules such as antibodies. 

These include ease of preparation, potentially lower antigenicity, and increased stability (Forssen 

and Willis 1998). RGD peptides have reported as effective in delivering cytotoxic molecules to the 

tumor vasculature (Arap, Pasqualini, and Ruoslahti 1998; Ellerby et al. 1999; Gerlag et al. 2001; 

Schiffelers et al. 2003). 

32.2 TUMOR VASCULATURE TARGETING 

There are many potential barriers to the effective delivery of chemotherapeutic agents, having 

narrowest therapeutic indices in all medicines to the disseminated malignant tumors. The nonselective 

toxic effects on normal tissues restrict the dose of the anti-cancer agents. Through selective 

delivery of chemotherapeutic agents into the malignant tumors, the concentration of drugs in the 

tumors could be increased. This results in a decrease in the amount and types of nonspecific 

toxicites and an increased amount of drug that can be effectively delivered to the tumor. Monoclonal 

antibodies have been used to provide existing targeting opportunities. However, this 

approach has met with limited success; only a few antigens are known, and their expression on 

the cells within an individual tumor is not necessarily uniform. Moreover, antibodies against tumor 

antigens poorly penetrate into solid tumors (Dvorak et al. 1991). Furthermore, antibody targeted 

drug delivery may be impaired by clonal selection for cells that have lost the tumor antigen because 

tumor cells are genetically unstable and adopt for mutations advantageous for growth. The targeting 

of drug delivery systems to the tumor vasculature overcomes some of the problems associated with 

conventional tumor targeting. 

Angiogenesis is an important process in tumor growth as well as in metastasis. Growth and 

survival of tumor cells depend on oxygen and nutrients supplied by the blood. As a consequence, 

the size of a tumor is restricted to a few cubic millimeters without the recruitment of new blood 

vessels. The architecture of a solid tumor is such that numerous layers of tumor cells are fed by one 

blood vessel. This poses a significant barrier for selective delivery of cytotoxic drugs into malignant 

tumors that need to extravasate from blood into the tumor tissue to reach the target cell; the larger 

the chosen carrier, the less accesible the tumor tissue will be. In contrast, endothelial cells lining the 

tumor vasculature are easily accessible for these macromolecular preparations. The tumor vasculature 

targeting is a valuable tool for selective delivery of drugs is which either do not reach the 

target cells insufficiently or are too toxic to non-target cells when administered systemically. 

Figure 32.1 shows the principle of targeted chemotherapy to tumor blood vessels. 

32.2.1 RATIONALE FOR TUMOR VASCULATURE TARGETING 


Vascular targeting offers several benefits over conventional targeting to tumor cells (Table 32.1). 

The endothelium of the tumor vasculature is readily accessible to a circulating probe, whereas a 

conventional tumor targeting probe has to overcome the high interstitial pressure within the solid 


RGD-Modified Liposomes for Tumor Targeting 645 

Receptors overexpressed 

on tumor 

vasculature 

Carrier constructs 

bearing antibody or 

ligands 

Blood vessel 

Tumor 

tissue 

FIGURE 32.1 Principle of targeted chemotherapy to tumor blood vessels. 

No 

penetration 

Normal tissue 

Peptide 

Liposome 

tumor by penetrating among closely packed tumor cells and dense tumor stroma (Dvorak et al. 

1991). 

The survival and growth of the tumor cells largely depend on blood supply to them. An anticancer 

therapy directed against the tumor vasculature does not have to eliminate every endothelial 

× 

Polyethylene glycol chain 

TABLE 32.1 

Comparison Between Conventional Tumor Targeting and Vascular Tumor Targeting 

Feature Conventional Tumor Targeting Vascular Tumor Targeting 

Pharmacokinetic system Multiple compartment Single compartment 

Primary target cell type Malignant tumor cells Non-malignant activated tumor 

vascular cells 

Target cell accessibility Poor Good 

Receptor Tumor antigen Angiogenic markers 

Receptor expression Uneven and heterogeneous Possibly uniform and homogeneous 

Homing moiety Monoclonal antibodies Monoclonal antibodies, peptides 

Therapeutic moiety Cytotoxic drug Cytotoxic drug, angiogenesis 

inhibitors 

Target cell genome Genetically unstable cells Genetically stable, diploid cells 

Acquired cytotoxic drug resistance Yes Not yet reported 

Posttargeting amplification loop Absent Present 

Complete elimination of target cell Required Not required 

Potential clinical applicability Limited by antigen heterogeneity Broad, as angiogenic markers may be 

shared 


646 

cell; partial denuding of the endothelium is likely to lead to the formation of an occlusive thrombus 

that stops flow to the part of the tumor served by the vessel. In addition, tumor vasculature targeting 

has an intrinsic amplification mechanism. It has been estimated that elimination of a single endothelial 

cell can inhibit the growth of approximately 100 tumor cells (Denekamp 1993). Moreover, 

tumor endothelial cells are diploid and non-malignant, and they are unlikely to lose cell surface 

target receptors or acquire drug resistance through mutation and clonal evolution. 

Oncologists have long recognized that tumors commonly develop resistance to chemotherapy, 

whereas normal tissues do not. Therefore, toxicity to normal tissues such as chemotherapy-induced 

myelosuppression continues to occur even after tumor cells have become drug resistant and 

progress in the face of continued treatment. Endothelial cells, being non-malignant cells, are 

expected to behave in a manner analogous to bone marrow cells. Long-term antiangiogenic 

therapy has not been shown to produce drug resistance in experimental animals (Boehm et al. 

1997) or in clinical trials (Folkman 1997). 

32.2.2 MOLECULAR TARGETS IN ANGIOGENIC TUMOR VASCULATURE 

For successful targeting to angiogenesis-associated blood vessels, angiogenic endothelial cells need 

to be discriminated from the normal quiescent endothelium. In the past decade, using cellular and 

molecular biological approaches, many receptors, e.g., a vb 3 (vitronectin receptor), VEGFR, MMP- 

2/-9, endoglin (CD105), and aminopeptidase N (CD13), have been identified to be differentially 

upregulated on tumor endothelial cells. These receptors can be explored as target determinants for 

drug targeting purposes (Table 32.2). 

Several approaches exploiting different target epitopes/determinants were investigated for their 

potential to interfere with the endothelial neovascularization. VEGFR-I and VEGFR-2 are overexpressed 

on tumor vasculature, while being present at a low density in the surrounding normal 

tissues. VEGF-dephtheria toxin conjugate treatment of tumor bearing mice resulted in selective 

vascular damage in the tumor tissue and inhibited tumor growth (Olson et al. 1997). Histological 

studies revealed that conjugate treatment spared the blood vessels of normal tissues such as liver, 

lung, and kidney from being damaged. 

Endoglin (CDI05) is a transmembrane glycoprotein, over-expressed in the vasculature of 

tumors and other tissues undergoing vascular remodelling (Miller et al. 1999). Differential 

TABLE 32.2 

Potential Target Epitopes on Tumor Vascular Endothelium 

Target Epitope Reference 

30.5-kDa antigen Hagemeier et al. (1986) 

CD34 Schlingemann et al. (1990) 

VEGF/VEGF receptor complex Ramakrishnan et al. (1996) 

VEGF receptor Dvorak et al. (1991) 

Endosialin Rettig et al. (1992) 

E-selectin Nguyen et al. (1993) 

aV integrins Brooks et al. (1994) 

Endoglin Burrows et al. (1995) 

Tie-2 Sato et al. (1995) 

TNF a receptor Eggermont et al. (1996) 

CD44 Griffioen et al. (1997) 

Angiostatin receptor Moser et al. (1999) 

Endostatin receptor Chang et al. (1999) 

MMP-2/MMP-9 Koivunen et al. (1999) 




up-regulation of endoglin presents an interesting opportunity to the selective delivery of cytotoxic 

molecules to the target endothelial cells. Several monoclonal antibodies specific for human endoglin 

have been produced (Seon et al. 1997). Radioiodinated monoclonal antibodies (10 mi) given to 

tumor-bearing animals significantly inhibited the tumor growth. This indicates the clinical potential 

of cytotoxic agent targeting using endoglin (Tabata et al. 1999). 

Other receptors over-expressed in tumor vessels include FGF receptors and Tie-2 receptors. 

Davol and colleagues prepared an endothelial cell-specific cytotoxic conjugate (Davol et al. 1995) 

by chemically linking the plant-derived ribosomal inhibitory protein saporin to FGF. The FGFsaporin 

conjugate inhibited proliferation of endothelial cells effectively in vitro. Radioimmunotherapy 

using 213 Bi or 131 I targeted to lung vasculature with a monoclonal antibody against 

thrombomodulin has been reported (Kennel and Mirzadeh 1998). Because thrombomodulin is 

selectively expressed in the pulmonary blood vessels, the antibody homed to lungs resulted in 

destruction of small tumor colonies. Cationic liposomes were used to target angiogenic vasculature 

in mouse pancreatic islet cell tumors. Confocal microscopy demonstrated a 15- to 33-fold preferentially 

higher uptake of the liposomes by angiogenic than normal endothelial cells (Thurston et al. 

1998). 

Erkki Ruoslahti and colleagues (1996) developed a novel targeting strategy by using polypeptidal 

system capable of delivering cytotoxic drugs selectively to integrins. An in vivo selection 

of phage display libraries identified peptides that specifically home to tumor blood vessels. 

Ruoslahti’s research group identified two major classes of peptides, one containing the RGD 

motif and the other containing an NGR motif. These polypeptides were then chemically linked 

to the anti-cancer drug doxorubicin. Treatment of breast carcinoma-bearing mice with the conjugated 

doxorubicin caused selective vascular damage in the tumors with consequential strong antitumor 

effect at a 10–40 times lower concentration as compared of free doxorubicin, whereas liver 

and heart toxicity was significantly reduced compared to that observed with free doxorubicin (Arap, 

Pasqualini, and Ruoslahti 1998). Their results illustrate the potential of targeting therapeutic agents 

to integrins expressed on the vasculature of tumors as an effective means of cancer treatment. Other 

approaches have also been used, e.g., toxic drugs (Arap, Pasqualini, and Ruoslahti 1998; Arora 

et al. 1999), apoptosis-inducing agents (Ellerby et al. 1999), cytokines (Curnis et al. 2000), and 

therapeutic genes (Hood et al. 2002) were delivered to the tumor endothelial cells, leading to 

reduced tumor growth. Furthermore, angiogenesis and tumor growth were inhibited upon occlusion 

of blood vessels by targeting coagulation factors to tumor vasculature (Huang et al. 1997; Ran et al. 

1998). Recently, immune effector cells have successfully been targeted to angiogenic blood vessels, 

with resultant lysis of tumor endothelial cells and suppression of tumor growth (Niederman et al. 

2002). 

32.3 ROLE OF RGD IN INTEGRIN- AND FIBRONECTIN-MEDIATED BIOEVENTS 

Among the many molecules that are selectively expressed in angiogenic vasculature, the integrins 

show particular promise in tumor targeting. Saiki and co-workers reported that tumor angiogenesis 

can be inhibited by blocking the interaction between integrins and the RGD motif-containing extra 

cellular matrix proteins (Saiki et al. 1990). One of the most studied target epitopes on angiogenic 

neovasculature is the avb3 integrin. Studies show that avb3 and avb5 integrins are upregulated in 

angiogenic endothelial cells (Brooks, Clark, and Cheresh, 1994a) and the inhibition of these 

integrins by antibodies, cyclic RGD peptides (Brooks et al. 1994b) and RGD peptidomimetics 

(Carron et al. 1998) can block neovascularization. 

32.3.1 INTEGRINS 

Members of the integrin family of adhesion molecules are non-covalently-associated a/b heterodimers 

that mediate cell–cell, cell–extracellular matrix, and cell–pathogen interactions by binding 


648 

to distinct, but often overlapping, combinations of ligands. Eighteen different integrin a-subunits 

and eight different b-subunits have been identified in vertebrates that form at least 24 a/b heterodimers, 

perhaps making integrins the most structurally and functionally diverse family of celladhesion 

molecules (Hynes 1992; Springer 1994) (Figure 32.2). Half of integrin a-subunits 

contains inserted (I) domains that are the principal ligand-binding domains (Shimaoka 2002). 

The complexity and structural and functional diversity of integrins allow this family of adhesion 

molecules to play a pivotal role in broad contexts of biology, including inflammation, innate and 

antigen specific immunity, haemostasis, wound healing, tissue morphogenesis, and regulation of 

cellular growth and differentiation. Conversely, disregulation of integrins is involved in the pathogenesis 

of many diseased states from autoimmunity to thrombotic vascular diseases to cancer 

metastasis (Curley, Blum, and Humphries 1999). Therefore, extensive efforts have been directed 

toward the discovery and development of integrins antagonists for clinical applications. 

Significant advances have been made in targeting aIIbb3 integrins on platelets for inhibiting 

thrombosis (Scarborough and Gretler 2000), avb3 and avb5 for blocking tumor metastasis, angiogenesis 

and bone resorption (Varner and Cheresh 1996) and b2 integrins and a4 integrins on 

leukocytes for treating autoimmune diseases and other inflammatory disorders (Giblin and Kelly 

2001; Yusuf-Makagiansar et al. 2002). Small-molecules, i.e., integrin inhibitors not only interfere 

with ligand binding but also stabilize particular integrin conformations that have provided insights 

into integrin structural rearrangements. 

The general structural features of all integrins appear to be similar. Both a- and b-subunits are 

transmembrane glycoproteins with large globular amino-terminal extracellular domains that, 

together, make up an ellipsoidal head (Figure 32.3). Each subunit provides a relatively thin leg 

that traverses across the plasma membrane and ends into a relatively small cytoplasmic tail of less 

than 60 amino acids. The only known integrin that does not fit this general description is b4 integrin 

that has a cytoplasmic domain of close to 1000 amino acids (Suzuki and Naitoh 1990; Tamura et al. 

1990). 

αL 

αM 

αX 

αD 

αE β7 

β4 

α1 

α1 

0 

α1 

1 

α2 

α3 

α4 

α5 

α6 

α7 

α8 

α9 

FIGURE 32.2 Integrin heterodimer composition. Integrin a- and b-subunits form 24 heterodimers that recognize 

distinct, but overlapping, ligands. (Adapted from Hynes, R. O., Cell, 69, 11, 1992.) 


β1 


αv 

β3 

β5 

β6 

β8 

αII 

b


26−28 nm 

4 nm 

8 nm 

12−14 nm 

S - S 

2 nm 

a b 

6 nm 

-S-S- 

The b1 integrins that can bind fibronectin include a3b1, a4b1, a5b1, and avb1. Many integrins, 

including the a3b1, a5b1, avb1, avb3, avb5, avb6, and the aIIbb3, recognize the RGD site located in 

adhesive proteins (Table 32.3). The fibronectin-specific integrin that consists of an a5 subunit and a 

b 1 subunit is the major fibronectin receptor expressed on most of the cells. This integrin mediates 

such cellular responses to fibronectin as adhesion, migration, assembly of a cytoskeleton, and 

assembly of the fibronectin extracellular matrix (The a 5b 1 integrin interacts with the central cell 

adhesive region of fibronectin and requires both the RGD and synergy sites for maximal binding 

(Obara and Yoshizato 1995). The major platelet integrin a IIhb 3 also recognizes a similar synergy 

site in fibronectin for mediating platelet interactions with fibronectin (Bowditch et al. 1994). 

Ca 2+ 

Ca 2+ 

Ca 2+ 

Ca 2+ 

-S-S- 

-S-S- 

-S-S- 

-S-S- 

-S-S- 

-S-S- 

-S-S- 

-S-S- 

-S-S- 

FIGURE 32.3 The structural model of a5b1 integrin adhesion molecule. A structural model based on 

predicted primary structure and electron microscopy visualization of purified integrin is shown. The a5subunit 

consists of part of the globular extracellular domain and one leg that contains predicted 12 b strands 

(gray box). The b1- subunit makes up the remainder of the globular extracellular domain. The b1- subunit also 

contains five cysteine-rich repeats. Both the a5- and b1- subunits have transmembrane domains and small 

cytoplasmic domains. (Adapted from Varner, J. A. and Cheresh, D. A., Important Adv. Oncol., 69, 1996.) 


650 

TABLE 32.3 

Integrin Receptors and Their Binding Sites 

Receptor Ligand 

32.3.2 FIBRONECTIN 

Receptor-Mediated 

Action 


Amino Acid Sequence 

Recognized 

a2b1 Collagen Adhesion DGEA 

a5b1 Fibronectin Adhesion RGD 

a 6b 1 Laminin Adhesion YIGSR 

aIIbb3 Fibrinogen Aggregation KQAGDV or RGD 

Fibronectin Aggregation RGD 

von Willebrand factor Aggregation RGD 

Vitronectin Aggregation RGD 

a6b3 Vitronectin Adhesion RGD 

Fibrinogen Adhesion RGD 

Fibronectinvon Adhesion RGD 

von Willebrand factor Adhesion RGD 

Fibronectin is a large adhesive glycoprotein found in extracellular matrices and body fluids 

(Mosher 1989; Carsons 1990; Hynes 1990). The primary structure of fibronectin is comprised of 

three different types of homologous repeating units or modules (Figure 32.4). There are several 

alternatively spliced forms of fibronectin that result from the deletion or insertion of complete type 

III modules. In addition, one particular region designated as IIICS can be partially inserted or 

deleted in some isoforms by alternative splicing. The homologous modules comprising of fibronectin 

are arranged in protease-resistant domains that are separated by more flexible, proteasesusceptible 

regions. When cleaved from intact fibronectin by partial proteolysis or expressed in 

bacterial or mammalian cells and individually purified, these domains often retain the specific 

binding functions of intact fibronectin such as those for he-parin, fibrin, denatured collagen 

(gelatin), and cell surface receptors. 

Fibronectin contains at least two distinct regions that can independently interact with distinct 

cell surface receptors. The first fibronectin cell-adhesive site to be identified was isolated in the 

form of protease-resistant fragments of 110–120 kDa, 75 kDa, and 37 kDa (Ruoslahti 1981; 

Hayashi and Yamada 1983; Zardi et al. 1985; Nagai et al. 1991) derived from the central 

portion of the protein. Such fragments of fibronectin retained similar cell adhesive activities as 

those of intact fibronectin. The cell adhesive activity attributed to these fragments was initially 

localized to the tenth type III module in the form of an 11.5 kDa pepsin fragment (Pierschbacher, 

Heparin Collagen 

ED-B Cell 

Heparin Fibrin 

ED-A III-C-S 

H2NI II III II II II 

I 

II 

I 

II 

I 

II 

I 

II II 

I I 

III 

II II II II II II 

I I I I I 

PHSRN RGD 

II 

I 

II 

I 

II 

I 

LDV 

II 

I 

REDV 

I I COOH 

S S 

S S 

FIGURE 32.4 Structure of fibronectin. Fibronectin is composed of three types of internal repeating modules 

designated as Type I, Type II, and Type II. The ED-A, ED-B, and III-C-S modeules can be present or absent in 

some forms of fibronectin as a result of alternative splicing. The binding domains are indicated at the top. The 

central cell-binding domain consists of ninth and tenth Type III nodules, containing the minimal PHSRN and 

RGD cell recognition sequence (Adapted from Mosher, D. F., Fibronectin, Academic Press, NY, p.474, 1989.) 



Hayman, and Ruoslahti 1981) and to a smaller peptide with the sequence Gly-Arg-Gly-Asp-Ser 

(GRGDS). Although the 11.5 kDa fragment and synthetic peptides containing the RGD sequence 

can inhibit fibronectin cell-adhesive functions in vitro and in vivo when added as soluble inhibitors 

(Pierschbacher and Ruoslahti 1984; Lash, Linask, and Yamada 1987), they only poorly promote 

cell adhesion mediated by the major fibronectin receptor a5b1 integrin, and their affinities are too 

low to be estimated in direct binding studies (Akiyama and Yamada 1985), suggesting that 

sequences outside of the tenth type III module are also important for maximal cell binding 

and adhesion. 

This additional fibronectin sequence that is important for maximal cell adhesive activity was 

identified and characterized using a series of mutants of the central fibronectin cell adhesive region 

expressed in E. coli (Aota, Nagai, and Yamada 1991) and anti-fibronectin monoclonal antibodies 

(Nagai et al. 1991). Fragments containing the RGD sequence but truncated approximately 10– 

14 kDa to the amino-terminal side of the RGD sequence were !4% as active as compared to intact 

fibronectin in mediating cell adhesion, whereas fragments containing these amino terminal 

sequences retained O97% of the activity of intact fibronectin. The novel, amino-terminal (non- 

RGD) site appeared to act synergistically with the RGD site to promote better cell adhesion, leading 

to its designation as a synergistic adhesive site or synergy site. 

The biological function of the synergistic cell adhesive site was characterized by using a panel 

of anti-fibronectin monoclonal antibodies (mAbs) developed to bind a 37 kDa cell adhesive fibronectin 

fragment. One of these antibodies, designated 8E3, bound to the ninth type III module at a 

site approximately 14–16 kDa to the amino-terminal side of the RGD sequence and close to the 

synergy site identified by mutational studies. Other mAbs that inhibited cell adhesion such as 333 

(Akiyama et al. 1985) and 16G3 bound to the tenth type III module and inhibited the RGD site. 

Interestingly, there was also an antibody, designated 13G12, that bound to fibronectin between the 

RGD and synergy sites but did not inhibit cell adhesion. Antibodies that bound near the RGD and 

synergy sites could each individually inhibit cell spreading at high concentrations. Furthermore, 

mAb 8E3 and mAb 16G3, at concentrations too low to inhibit individually spreading, can be highly 

inhibitory in combination, underscoring the synergistic nature of the two cell adhesive sites. 

Antibody inhibition experiments have also shown that both the RGD and synergy sites function 

in cell migration and cytoskeleton assembly on fibronectin substrates as well as in the assembly of a 

fibronectin extracellular matrix (Nagai et al. 1991), indicating that both sites are required for a 

range of fibronectin activities. 

32.3.3 RGD-MEDIATED METASTASIS INHIBITION 

A major cause of morbidity and death as a result of cancer is the metastasis of cells from the primary 

tumor to distant sites where secondary tumors become established. As schematically shown in 

Figure 32.5, metastasis is a multistep process. These steps include detachment of cells from the 

tumor mass, degradation of basement membrane, migration to and invasion into the vascular or 

lymphatic systems, arrest at a distant site, adhesion to the vascular endothelium, degradation of 

basement membrane, extravasation, and migration to and proliferation at the secondary site. Many 

of these steps require cell adhesive interactions or loss of adhesion. Tumor cell adhesion to 

components of the extracellular matrix and basement membranes is mediated by specific cell 

surface receptors that bind to extracellular adhesive proteins. The fibronectin-integrin system has 

provided a valuable model system for the study of molecular mechanisms of ligand-receptor 

interactions involved in cell adhesive steps in metastasis. 

Several different assay systems have been used to directly analyze the role of fibronectin and 

integrins in in vivo metastasis, but they all fall into two broad classes. The experimental metastasis 

models mainly involve the intravenous injection of tumor cell suspensions into mice and subsequent 

quantitation of either the number of metastatic colonies or the sizes of colonies usually 

in the lungs or liver. Experimental metastasis models, however, exclude the earlier steps of the 


652 

Basement 

Membrane 

Carcinoma 

Invasion 

Sarcoma 

Capillary 

Migration 


Blood vessel or 

Lymphatic system 

Proliferation 

Extravasation 

Arrest 

FIGURE 32.5 Major steps in metastasis. Tumor cells are shown invading the circulatory system from a 

carcinoma by degrading and migrating through a basement membrane or from a sarcoma. Cells separate 

from the primary tumor, enter the vasculature, and eventually arrest in capillaries where they penetrate and 

migrate through basement membrane and underlying connective tissue to the metastatic site where colonization 

and proliferation occur. 

metastatic process involving detachment from the tumor mass and invasion into the vasculature 

and, subsequently, replicate only following the final steps of hematogenous metastasis. Spontaneous 

metastasis models where tumors or cells are implanted into mice and subsequent 

metastases to distant organs are quantitated either by counting colonies or measuring colony size 

are more complicated, but they are required to analyze of the earlier steps of the metastatic cascade. 

Synthetic peptides containing the Gly-Arg-Gly-Asp-Ser (GRGDS) sequence derived from 

fibronectin were assayed by examining their effects on lung colonization of B16-F10 murine 

melanoma cells in syngeneic C57BL/6 mice using an experimental metastasis model (Humphries, 

Olden, and Yamada 1986). GRGDS peptide that had a circulatory half-life of approximately eight 

minutes specifically inhibited lung colonization in a concentration-dependent manner. The major 

effect of the GRGDS peptide appeared to be to inhibit arrest of melanoma cells in the lungs without 

affecting the size of either melanoma cell clusters in suspension or lung colonies. The efficacy of 

peptide treatment was unaltered in animals with impaired platelet function and in animals lacking 

natural killer cells. Taken together, these results suggest that the GRGDS peptide could inhibit 

metastasis by disrupting an early adhesive process. As shown in Figure 32.6, a single administration 

of the GRGDS peptide also had the dramatic effect of increasing the survival of mice when 

co-injected with melanoma cells (Humphries, Yamada, and Olden 1988). 

RGD peptides and RGD mimetics are very promising molecules. RGD peptides not only inhibit 

metastatic colony formation but a single administration of the peptide can dramatically increase 

animal survival also however, the use of such agents have some drawbacks. One potential problem is 

the relatively high concentrations of peptide required for activity. Possible solutions to the problems 

may include the use of cyclic peptides and designer peptides that reportedly increased the affinity for 

cell surface integrins or the use of repeating polymers containing RGD units (Saiki et al. 1989; 

Murata et al. 1991; Yamamoto et al. 1994). Polymers containing RGD sequences have the dual 

advantages of augmenting the affinity of small peptide sequences through multivalent interactions 

and the possibility of solving the problem of the relatively short circulatory half-life of GRGDS 

peptides. An alternative approach would be to develop novel, small, synthetic integrin antagonists. 

Such compounds that inhibit RGD-dependent and LDV-dependent cell adhesion have already been 

synthesized and tested in mice (Greenspoon et al. 1993, 1994). 



Survival (%) 

100 

75 

50 

25 

0 

• 

• 

• 

Treated 

• 

• 

• 

• 

Untreated 

• 

• 

30 60 120 180 240 300 360 

Days post injection 

FIGURE 32.6 Effect of GRGDS peptide on survival of C57BL/6 mice injected with B16F10 murine melanoma 

cells. Mice were co-injected with 3!10 4 B16F10 melanoma cells mixed with 3 mg GRGDS paptide. 

(Data are from Humphries et al., J. Clin. Invest., 81, 782, 1988.) 

32.4 RGD-MODIFIED LIPOSOMES FOR CANCER THERAPEUTICS 

Cytotoxic drug incorporation into liposomes has been reported since the mid-1970s. These early 

reports laid the groundwork for selecting therapeutic agents amenable to incorporation by liposomes 

as well as determining optimal liposome size and net charge required for effective drug 

delivery (Gregoriadis 1976). Early reports also noted the problems associated with liposomemediated 

drug delivery; liposomes were removed from circulation by fixed macrophages in the 

RES, particularly in the liver and spleen. Although first-generation liposome-incorporated drugs 

may be effective in macrophage-related diseases, particularly in the liver and spleen, many other 

tumors would require other drug-targeting mechanisms. Later studies refined many of the considerations 

required for effective liposome-mediated drug delivery (Drummond et al. 1999). 

In 2005, Pattillo et al. developed RGD-modified immunoliposmes for targeting the antivascular 

drug combretastatin to irradiated mouse melanomas. Combretastatin was incorporated into liposomes 

with surfaces modified by the addition of cyclo(Arg-Gly-Asp-D-Phe-Cys) (RGD) to create 

an immunoliposome. In this transplanted tumor model (B16-F10 melanoma), there was no significant 

increase in the tumor volume when treated with a single dose of 5-Gy radiation and RGD 

modified immunoliposomes (14.5 mg/kg of combretastatin) during the initial six days post treatment; 

all other treatment groups exhibited exponential growth curves after three days. The 

treatment resulted in a 5.1-day tumor growth delay compared to untreated controls. These findings 

indicate that preferential targeting of antivascular drugs to irradiated tumors results in significant 

tumor growth delay. 

Various strategies are adopted to target the sites of tumor-associated angiogenesis or other 

inflammatory environments with liposome-immobilized ligands where adhesion receptors (or 

ligands) serve as molecular targets (Figure 32.7). Targeting can be achieved using RGD 

or YIGSR immobilized on liposome surface and interacting with integrin receptors on normal or 

malignant cells that over-express them. Targeting can be achieved using synthetic antibodies 

anchored on the liposomes and bind integrins and other cell adhesion molecules molecules by 

mimicking their natural ligands. Targeting can be achieved by immobilizing specific carbohydrate 

ligands (sLe x and sLe a ) on liposomes that show specific interaction with selectin. These strategies 

are directed either to block the normal recruitment of leucocytes during vasculogenesis or for 

selective targeting of chemotherapeutic agents to tumor-associated vasculature. 

Cyclic RGD peptide, cyclo(Arg-Gly-Asp-Phy-Lys) anchored sterically stabilized liposomes 

(RGD-SL) were investigated for selective and preferential presentation of carrier contents at angiogenic 

endothelial cells over-expressing alphavbeta3 integrins on and around tumor tissue and for 

assessing their targetabilty (Dubey et al. 2004; Dubey 2005). Liposomes were prepared using 


654 

YIGSR 

RGD 

distearoylphosphatidylcholine (DSPC), cholesterol, and distearoylphosphatidylethanolamine-polyethyleneglycol-RGD 

peptide conjugate (DSPE-PEG-RGD) in a molar ratio 56:39:5. The control 

RAD peptide anchored sterically stabilized liposomes (RAD-SL), and liposome with 5 mol% PEG 

(SL) without peptide conjugate that had similar lipid composition were used for comparison. The 

average size of all liposome preparations prepared was approximately 105 nm and maximum drug 

entrapment was 10.5C/K1.1%. In vitro endothelial cell binding of liposomes exhibited 7-fold 

higher binding of RGD-SL to HUVEC in comparison to the SL and RAD-SL (Figure 32.8a). 

Table 32.4 and Figure 32.9 indicate RGD-modified liposomes mediated spontaneous lung 

metastasis and angiogenesis assays, respectively. RGD peptide-anchored liposomes are significantly 

(p!0.01) effective in the prevention of lung metastasis and angiogenesis compared to 

sLe x 

sLe a 

(a) (b) 

FIGURE 32.7 Liposome-immobilized ligands to target the cell adhesion molecules on tumor vasculature and 

extracellular matrix. These liposomes serve as artificial leucocyte to mediate interactions with integrins (a), 

integrins and other CAM molecules (b), and selectins (c) with the help of immobilized ligands or anti-receptor 

antibody. (Adapted from Vyas, S. P., and Khar, R. K., Novel Carrier Systems, CBS publishers, New Delhi, 

p.535, 2001.) 

% of maximum HUVEC binding 

120 

100 

80 

60 

40 

20 

0 0 

SL RAD-SL RGD-SL RGD-SL + RGD-SL + 

Control Free 5-FU SL RGD-SL RAD-SL 

(a) 

free RGD 

Formulation 

free RAD 

(b) 

Formulation 

FIGURE 32.8 (a) Binding of RGD-modified liposomes with HUVEC cells. (b) Effect of various formulations 

on tumor angiogenesis by an intradermal injection of B16F10 melanoma. Results are given as means (S.D.) 

NZ5(P!0.01). 5-FU, SL, RGD-SL, and RAD-SL indicate free 5-fluorouracil, stealth liposomes (SL), RGDanchored 

stealth liposomes (RGD-SL), and RAD-anchored stealth liposomes (RAD-SL), respectively 

(Adapted from Dubey, P. K., et al., J. Drug. Target., 12, 257, 2004). 


Angiogenesis (No. of vessels) 

(c) 

30 

25 

20 

15 

10 

5 



TABLE 32.4 

Effect of Free 5-FU, SL, RGD-SL, and RAD-SL on Spontaneous Lung Metastasis in BALB/c 

Mice by an Intra-Footpad Injection of B16F10 Melanoma 

Formulations 

Average Number 

of Metastasis/ 

Mouse 

Average Number of Metastasis/ 

Tumor Bearing Mouse 

Number of Metastasis Free Mice 

(NZ10) 

Control 52G15.5 52G15.5 0 

Free 5-FU 46G13.6 46G13.6 0 

SL 26G8.6 32G7.65 2 

RGD-SL 8G4.2* 16G6.58 5 

RAD-SL 28G6.5 31G8.7 1 

Source: Dubey et al., J. Drug. Target.,12, 2004. *P!0.01 

free 5-FU, SL, and RAD-SL. In therapeutic experiments, 5-FU, SL, RGD-SL, and RAD-SL were 

intravenously administered on day 4 at the dose of 10 mg 5-FU/kg body weight to B16F10 tumor 

bearing BALB/c mice, resulting in effective regression of tumors compared with free 5-FU, SL, and 

RAD-SL (Dubey et al. 2004). Results indicated that cyclic RGD peptide anchored sterically 

stabilized liposomes bearing 5-FU were more significantly (p!0.01) active against primary 

tumor and metastasis than the non-targeted sterically stabilized liposomes and free drug 

(Figure 32.9). Vascular damage and reduced vascularization were recorded in tumors treated 

with RGD-SL, but not in the tumor treated with SL (Figure 32.10). These results also confirmed 

that RGD peptide-anchored liposomes cause a marked improvement in therapeutic efficacy and 

growth inhibitory effect on tumors. 

RGD-modified sterically stabilized liposomes have also been evaluated to improve the antitumor 

efficacy of doxorubicin (Schiffelers et al. 2003; Xiong et al. 2005). RGD peptide was coupled 

to the distal end of the poly (ethylene glycol)-coated liposomes. RGD peptide anchored sterically 


2000 

1800 

1600 

1400 

1200 

1000 

800 

600 

400 

200 

0 

0 

4 6 8 10 12 14 16 18 20 

Days following tumor implantation 

FIGURE 32.9 Anti-tumor efficacy of control (!), free 5-FU (%), SL (&), RAD-SL (C), and RGD-SL (:)in 

mice bearing B16F10 melanoma subcutaneously inoculated in the back. The arrow represents the time of 

injection. Results are given as means (S.D.) NZ10 (P!0.01). (Adapted from Dubey, P.K. et al., J. Drug. 

Target., 12, 257, 2004.) 


656 

FIGURE 32.10 (See color insert following page 522.) Histopathological analysis (hematoxylin and eosin stain) 

of B16F10 tumors treated with (a) Stealth liposomes and (b) RGD-modified stealth liposomes (250X). (Adapted 

from Dubey, P. K., Development and characterization of biochemical based strategies for tumor targeting. Ph. D. 

thesis, Department of Pharmaceutical Sciences, Dr. H. S. Gour University, Sagar (MP) India, 2005.) 

stabilized liposomes (dose 5 mg doxorubicin/kg) have demonstrated prolonged circulation time and 

increased tumor accumulation and effective retardation in tumor growth compared with sterically 

stabilized liposomes without RGD modification. 

Holig and coworkers have isolated from phage display RGD motif libraries with novel highaffinity 

cyclic RGD peptides on the basis of their selectivity towards endothelial or melanoma cells 

(Holig et al. 2004). Although the starting sequences contained only two cysteine residues flanking 

the RGD motif, several of the isolated peptides possessed four cysteine residues. A high-affinity 

peptide (RGD10) (DGARYCRGDCFDG) constrained by only one disulfide bond was used to 

generate novel lipopeptides composed of a lipid anchor, a short flexible spacer, and the peptide 

ligand conjugated to the spacer end. Incorporation of RGD10 lipopeptides into liposomes resulted 

in specific and efficient binding of the liposomes to integrin-expressing cells. In vivo experiments 

applying doxorubicin-loaded RGD10 liposomes in a C26 colon carcinoma mouse model demonstrated 

improved efficacy compared with free doxorubicin and untargeted liposomes. 

Administration of large amounts of synthetic peptides based on the Arg-Gly-Asp (RGD) 

sequence has been shown to suppress tumor metastasis. To overcome the rapid degradation of 

peptides in the circulation, an RGD mimetic, L-arginyl-6-aminohexanoic acid (NOK), was synthesized 

and conjugated with phosphatidylethanolamine (PE) (NOK-PE) for liposomalization 

(Kurohane, Namba, and Oku 2000). Cell adhesion assays revealed that B16BL6 murine melanoma 

cells adhered to immobilized NOK-PE. This adhesion was inhibited by the addition of either 

soluble RGDS or NOK at similar concentrations in a dose-dependent manner. Administration of 

NOK-PE liposomes (equivalent to approximately 500 microg RGD peptides) via the tail vein 

completely inhibited lung colonization of B 16BL6 cells. The same dose of soluble NOK was 

not effective in inhibition of the tumor metastasis. In addition, injection of NOK-PE liposomes via 

the tail vein inhibited spontaneous lung metastasis of B16BL6 cells from the primary tumor site in 

the hind footpad. These results suggest that NOK, a structural mimetic of RGD, is capable of 

suppressing metastasis by blocking the binding of the integrins present on tumor cells to the RGDcontaining 

extracellular matrix. 

32.5 RGD-MODIFIED LIPOSOMES IN CANCER GENE THERAPY 


Gene therapy, by definition, aims at modifying the genetic program of a cell toward a therapeutic or 

prophylactic goal. Several gene therapy strategies for tumors are currently under evaluation and 

some of them include: the modification of the function of oncogenes and tumor suppressor genes; 



the modification of the host immune response toward the tumor; the disruption of the tumor 

neovascularization; the lysis of tumor cells with replication-competent viruses, and suicide gene 

therapy where an inactive prodrug is converted into a cytotoxic drug by gene-expressed enzymes. 

Both viral and non-viral gene vectors have been investigated and well documented (Ledley 1995). 

However, the inefficiency of current gene vectors in infecting targeted cells and their inability to 

selectively access the diseased cells systemically distributed are two major impediments that have 

to be overcome for further successful clinical applications. 

RGD peptides have been used to taget the lipid-protamine-DNA (LPD) lipopolyplexes to tumor 

cells (MDA-MB-231), expressing appropriate integrin receptors (Harvie et al. 2003). The incorporation 

of pegylated lipid into Lipid-Protamine-DNA (LPD-PEG) lipopolyplexes causes a 

decrease of their in vitro transfection activity. This can be partially attributed to a reduction in 

particle binding to cells. To restore particle binding and specifically target LPD formulations to 

tumor cells, the lipid-peptide conjugate DSPE-PEG5K-succinyl-ACDCRGDCFCG-COOH 

(DSPE-PEG5K-RGD-4C) was generated and incorporated into LPD formulations (LPD-PEG- 

RGD). LPD-PEG-RGD was characterized with respect to its biophysical and biological properties. 

The incorporation of DSPE-PEG5K-RGD-4C ligands into LPD formulations results in a 5- and 

15-fold increase in the LPD-PEG-RGD binding and uptake, respectively, over an LPD-PEG formulation. 

Enhancement of binding and uptake resulted in a 100-fold enhancement of transfection 

activity. Moreover, this transfection enhancement was specific to cells expressing appropriate 

integrin receptors (MDA-MB-231). Huh7 cells, known for their low level of avb 3 and a vb 5 integrin 

expression, failed to show RGD-mediated transfection enhancement. This transfection enhancement 

can be abolished in a competitive manner using free RGD peptide, but not an RGE control 

peptide. Results demonstrated RGD-mediated enhanced LPD-PEG cell binding and transfection in 

cells expressing the integrin receptor. These formulations provide the basis for effective, targeted, 

systemic gene delivery. 

In 2002, Fahr et al. developed a novel liposomal vector (Artificial Virus Particles; AVPs) for 

cancer gene therapy. Artificial virus-like particles (AVPs) represent a novel type of liposomal 

vector, resembling retroviral envelopes. AVPs are serum-resistant and non-toxic and can be 

endowed with a peptide ligand as a targeting device. These workers investigated if AVPs carrying 

cyclic peptides with an RGD integrin-binding motif (RGD-AVPs) were suitable for the specific and 

efficient transduction of human melanoma cells. In the experimentation, they prepared Plasmid 

DNA-low molecular weight non-linear polyethyleneimine complexes and packaged into anionic 

liposomes. Transduction efficiencies were determined after transient transfection of different cell 

lines in serum-free medium using green fluorescent protein or luciferase reporter genes. It was 

found that RGD-AVPs transduced human melanoma cells with high efficiencies of O60%. Efficient 

transduction was clearly dependent on the presence of the cyclic RGD ligand and was 

selective for melanoma cells. Using the melanocyte-specific tyrosinase promoter to drive transgene 

expression, the specificity of the vector system could be further enhanced. 

REFERENCES 

Akiyama, S. K. and Yamada, K. M., Synthetic peptides competitively inhibit both direct binding to fibroblasts 

and functional biological assays for the purified cell-binding domain of fibronectin, J. Biol. Chem., 

260, 10402, 1985. 

Akiyama, S. K. et al., The interaction of fibronectin fragments with fibroblastic cells, J. Biol. Chem., 260, 

13256, 1985. 

Aota, S., Nagai, T., and Yamada, K. M., Characterization of regions of fibronectin besides the arginineglycine-aspartic 

acid sequence required for adhesive function of the cell-binding domain using sitedirected 

mutagenesis, J. Biol. Chem., 266, 15938, 1991. 

Arap, W., Pasqualini, R., and Ruoslahti, E., Cancer treatment by targeted drug delivery to tumor vasculature in 

a mouse model, Science, 279, 377, 1998. 


658 


Arora, N. et al., Vascular endothelial growth factor chimeric toxin is highly active against endothelial cells, 

Cancer Res., 59, 183, 1999. 

Boehm, T. et al., Antiangiogenic therapy of experimental cancer does not induce acquired drug resistance, 

Nature, 390, 404, 1997. 

Bowditch, R. D. et al., Identification of a novel integrin-binding site in fibronectin: differential utilization by b, 

integrins, J. Biol. Chem., 269, 10856, 1994. 

Brooks, P. C., Clark, R. A., and Cheresh, D. A., Requirement of vascular integrin avb3 for angiogenesis, 

Science, 264, 569, 1994a. 

Brooks, P. C. et al., Integrin Alpha-v Beta-3 antagonists promote tumor regression by inducing apoptosis of 

angiogenic blood vessels, Cell, 79, 1157, 1994b. 

Brown, P. M. and Silvius, J. R., Mechanisms of delivery of liposome-encapsulated cytosine arabinoside to 

CV-1 cells in vitro: Fluorescence-microscopic and cytotoxicity studies, Biochim. Biophys. Acta, 1023, 

341, 1990. 

Burrows, F. J. et al., Endoglin is an endothelial cell proliferation marker that is selectively expressed in tumor 

vasculature, Clin. Cancer Res., 1, 1623, 1995. 

Carron, C. P. et al., A. peptidomimetic antagonist of the intergin alpha (V) beta 3 inhibits Leydig cell tunour 

growth and the development of hypercalcemia of malignancy, Cancer Res., 58, 1930, 1998. 

Carsons, S. E., Fibronectin in Health and Disease, CRC Press, Boca Raton, 1990. 

Chang, Z., Choon, A., and Friedl, A., Endostatin binds to blood vessels in situ independent of heparan sulfate 

and does not compete for fibroblast growth factor-2 binding, Am. J. Pathol., 155, 71, 1999. 

Curley, G. P., Blum, H., and Humphries, M. J., Integrin antagonists, Cell. Mol. Life Sci., 56, 427, 1999. 

Curnis, F. et al., Enhancement of tumor neaosis factor alpha antitumour immunothecapetic properties by 

targeted delivery to aminopeptidase N (CD13), Nat. Biotechnol., 18, 1185, 2000. 

Davol, P. et al., Saporin toxins directed to basic fibroblast growth factor receptors effectively target human 

ovarian teratocarcinoma in an animal model, Cancer, 76, 79, 1995. 

Denekamp, J., Angiogenesis, neovascular proliferation and vascular pathophysiology as targets for cancer 

therapy, Br. J. Radiol., 66, 181, 1993. 

Drummond, D. C. et al., Optimizing liposomes for delivery of chemotherapeutic agents to solid tumors, 

Pharmacol. Rev., 51, 691, 1999. 

Dubey, P. K., Development and characterization of biochemical based strategies for tumor targeting. Ph. D. 

thesis, Department of Pharmaceutical Sciences, Dr. H. S. Gour University, Sagar (MP) India, 2005. 

Dubey, P. K. et al., Liposomes modified with cyclic RGD peptide for tumor targeting, J. Drug Target, 12, 257, 

2004. 

Dvorak, H. F. et al., Distribution of vascular permeability factor (vascular endothelial growth factor) in tumors: 

Concentration in tumor blood vessels, J. Exp. Med., 174, 1275, 1991. 

Eggermont, A. M. et al., Isolated limb perfusion with high-dose tumor necrosis factor-alpha in combination 

with interferon-gamma and melphalan for nonresectable extremity soft tissue sarcomas: A multicenter 

trial, J. Clin. Oncol., 14, 2653, 1996. 

Ellerby, H. M. et al., Anticancer activity of targeted pro-apoptotic peptides, Natl. Med., 5, 1032, 1999. 

Fahr, A. et al., A new colloidal lipidic system for gene therapy, J. Liposome Res., 12, 37, 2002. 

Folkman, J., Angiogenesis and angiogenesis inhibition: An overview, EXS, 79, 1, 1997. 

Forssen, E. and Willis, W., Ligand targeted liposomes, Adv. Drug Del. Rev., 29, 249, 1998. 

Gabizon, A. and Papahadjopoulos, D., The role of surface charge and hydrophilic groups on liposome clearance 

in vivo, Biochim. Biophys. Acta, 1103, 94, 1992. 

Gerlag, D. M. et al., Suppression of murine collagen-induced arthritis by targeted apoptosis of synovial 

neovasculature, Arthritis Res., 3, 357, 2001. 

Giblin, P. A. and Kelly, T. A., Antagonists of b2 integrin-mediated cell adhesion, Ann. Rep. Med. Chem., 36, 

181, 2001. 

Greenspoon, N. et al., Structural analysis of integrin recognition and the inhibition of integrin-mediated cell 

function by novel nonpeptidic surrogates of the Arg-Gly-Asp sequence, Biochemistry, 32, 1001, 1993. 

Greenspoon, N. et al., Novel 0 F-S-CH, peptide-bond replacement and its utilization in the synthesis of 

nonpeptidic surrogates of the Leu-Asp-Val sequence that exhibit specific inhibitory activities on 

CD4 C T cell binding to fibronectin, Int. J. Pept. Res., 43, 417, 1994. 

Gregoriadis, G., The carrier potential of liposomes in biology and medicine (first of two parts), N. Engl. 

J. Med., 295, 704, 1976. 



Griffioen, A. W. et al., CD44 is an activation antigen on human endothelial cells, involvement in tumor 

angiogenesis, Blood, 90, 1150, 1997. 

Gyongyossy-Issa, M. I. C., Muller, W., and Devine, D., The covalent coupling of Arg-Gly-Asp-containing 

peptides to liposomes: Purification and biochemical function of lipopeptide, Arch. Biochem. Biophys., 

353, 101, 1998. 

Hagemeier, H. H. et al., A monoclonal antibody reacting with endothelial cells of budding vessels in tumors 

and inflammatory tissues, and non-reactive with normal adult tissues, Int. J. Cancer, 38, 481, 1986. 

Harvie, P. et al., Targeting of lipid-protamine-DNA (LPD) lipopolyplexes using RGD motifs, J. Liposome 

Res., 13, 231, 2003. 

Hayashi, M. and Yamada, K. M., Domain structure of the carboxy-terminal half of human plasma fibronectin, 

J. Biol. Chem., 258, 3332, 1983. 

Holig, P. et al., Novel RGD lipopeptides for the targeting of liposomes to integrin-expressing endothelial and 

melanoma cells, Protein Eng. Des. Sel., 17, 433, 2004. 

Hood, J. D. et al., Tumor regression by targeted gene delivery to the neovasculature, Science, 269, 2404, 2002. 

Huang, X. et al., Tumor infarction in mice by antibody-directed targeting of tissue factor to tumor vasculature, 

Science, 275, 547, 1997. 

Humphries, M. J., Olden, K., and Yamada, K. M., A synthetic peptide from fibronectin inhibits experimental 

metastasis of murine melanoma cells, Science, 233, 466, 1986. 

Humphries, M. J., Yamada, K. M., and Olden, K., Investigation of the biological effects of anti-cell adhesive 

synthetic peptides that inhibit experimental metastasis of B16-F10 murine melanoma cells, J. Clin. 

Investig., 81, 782, 1988. 

Hynes, R. O., Fibronectins, Springer-Verlag, NY, p. 546, 1990. 

Hynes, R. O., Integrins: Versatility, modulation, and signaling in cell adhesion, Cell, 69, 11, 1992. 

Kennel, S. J. and Mirzadeh, S., Vascular targeted radioimmunotherapy with 2I3 Bi-an-particle emitter, Nucl. 

Med. Biol., 25, 241, 1998. 

Kirpotin, D. et al., Sterically stabilized anti-HER2 liposomes: Design and targeting to human breast cancer 

cells in vitro, Biochemistry, 36, 66, 1997. 

Koivunen, E. et al., Tumor targeting with a selective gelatinase inhibitor, Natl. Biotechnol., 17, 768, 1999. 

Kurohane, K., Namba, Y., and Oku, N., Liposomes modified with a synthetic Arg-Gly-Asp mimetic inhibit 

lung metastasis of B16BL6 melanoma cells, Life Sci., 68, 273, 2000. 

Lash, J. W., Linask, K. K., and Yamada, K. M., Synthetic peptides that mimic the adhesive recognition signal 

of fibronectin: Differential effects on cell–cell and cell–substratum adhesionin embryonic chick cells, 

Dev. Biol., 123, 411, 1987. 

Ledley, F. D., Nonviral gene therapy: the promise of genes as pharmaceutical products, Hum. Gene. Ther., 6, 

1129,1995. 

Lopes de Menezes, D. E., Pilarski, L. M., and Allen, T. M., In vitro and in vivo targeting of immunoliposomal 

doxorubicin to human B-cell lymphoma, Cancer Res., 58, 3320, 1998. 

Lundberg, B., Hong, K., and Papahadjopoulos, D., Conjugation of apolipoprotein B with liposome and 

targeting of cells in culture, Biochim. Biophys. Acta, 1149, 305, 1993. 

Martin, F. J., Clinical pharmacology and antitumor efficacy of DOXIL (pegylated liposomal doxorubicin), In 

Medical Applications of Liposomes, Lasic, D. D. and Papahadjopoulos, D., Eds., Elsevier Science BV, 

New York, p. 635, 1998. 

Maruyama, K. et al., Characterization of in vitro immunoliposome targeting to pulmonary endothelium, 

J. Pharm. Sci., 79, 978, 1990. 

Miller, D.W. et al., Elevated expression of endoglin, a component of the TGF-b-receptor complex, correlates 

with proliferation of tumor endothelial cells, Int. J. Cancer, 81,568,1999. 

Moser, T. L. et al., Angiostatin binds ATP synthase on the surface of human endothelial cells, Proc. Natl Acad. 

Sci. USA, 96, 2811, 1999. 

Mosher, D. F., Fibronectin, Academic Press, New York, 1989. 

Murata, J. et al., Molecular properties of poly(RGD) and its binding capacities to metastatic melanoma cells, 

Int. J. Pept. Prot. Res., 38, 212, 1991. 

Nishiya, T. and Sloan, S., Interaction of RGD liposomes with platelets, Biochem. Biophys. Res. Commun., 224, 

242, 1996. 

Nguyen, M. et al., A role for sialyl Lewis-X/A glycoconjugates in capillary morphogenesis, Nature, 365, 

267,1993. Erratum in: Nature, 366, 368, 1993. 


660 


Niederman, T.M. et al., Antitumor activity of cytoxic T lymphocytes engineered to target vascular endothelial 

growth factor receptors, Proc. Natl. Acad. Sci. USA, 99, 709, 2002. 

Obara, M. and Yoshizato, K., Possible involvement of the interaction of the a, subunit of the a, b, integrin with 

the synergistic region of the central cell-binding domain of fibronectin in cells to fibronectin binding, 

Exp. Cell Res., 216, 273, 1995. 

Olson, T. A. et al., Targeting the tumor vasculature: inhibition of tumor growth by a vascular endothelial 

growth factor toxin conjugate, Int. J. Cancer, 73, 865, 1997. 

Papahadjopoulos, D. and Gabizon, A. A., Sterically stabilized (stealth) liposomes: Pharmacological properties 

and drug carrying potentialin cancer, In Liposomes as Tools in Basic Research and Industry, Philippot, 

J. R. and Schuber, F., Eds., CRC Press, Boca Raton, p. 177, 1995. 

Papahadjopoulos, D. et al., Sterically stabilized liposomes: Improvements in pharmacokinetics and antitumor 

therapeutic efficacy, Proc. Natl Acad. Sci. USA, 88, 11460, 1991a. 

Papahadjopoulos, D. et al., Sterically stabilized liposomes: Improvements in pharmacokinetics and antitumor 

therapeutic efficacy, Proc. Natl Acad. Sci. USA, 88, 11460, 1991b. 

Park, J. W. et al., Development of anti-p185 HER2 immunoliposomes for cancer therapy, Proc. Natl Acad. Sci. 

USA, 92, 1327, 1995. 

Pattillo, C. B. et al., Targeting of the antivascular drug combretastatin to irradiated tumors results in tumor 

growth delay, Pharm. Res., 22, 1117, 2005. 

Pierschbacher, M. D. and Ruoslahti, E., Cell attachment activity of fibronectin can be duplicated by small 

synthetic fragments of the molecule, Nature, 309, 30, 1984. 

Pierschbacher, M. D. and Ruoslahti, E., Influence of stereochemistry of the sequence Arg-Gly- Asp-Xaa on 

binding specificity in cell adhesion, J. Biol. Chem., 262, 17294, 1987. 

Pierschbacher, M. D., Hayman, E. G., and Ruoslahti, E., Location of the cell attachment site in fibronectin with 

monoclonal antibodies and proteolytic fragments of the molecule, Cell, 26, 259, 1981. 

Ramakrishnan, S. et al., Vascular endothelial growth factor-toxin conjugate specifically inhibits KDR/Flk-1positive 

endothelial cell proliferation in vitro and angiogenesis in vivo, Cancer Res., 56, 1324, 1996. 

Ran, S. et al., Infarction of solid Hodgkin’s tumors in mice by antibody-directed targeting of tissue factor to 

tumor vasculature, Cancer Res., 58, 4646, 1998. 

Rettig, W. J. et al., Identification of endosialin, a cell surface glycoprotein of vascular endothelial cells in 

human cancer, Proc. Natl Acad. Sci. USA, 89, 10832, 1992. 

Ruoslahti, E., RGD and other recognition sequences for integrins, Annu. Rev., Cell Dev. Biol., 12, 697, 1996. 

Ruoslahti, E. et al., Alignment of biologically active domains in the fibronectin molecule, J. Biol. Chem., 256, 

7277, 1981. 

Saiki, I. et al., Inhibition of the metastasis of murine malignant melanoma by synthetic polymeric peptides 

containing core sequences of cell-adhesive molecules, Cancer Res., 49, 3815, 1989. 

Saiki, I. et al., Anti-metastatic and anti-invasive effects of polymeric Arg-Gly-Asp (RGD) peptide, poly 

(RGD), and its analogues, Jpn. J. Cancer Res., 81, 660, 1990. 

Sato, T.N. et al., Distinct roles of the receptor tyrosine kinases Tie-1 and Tie-2 in blood vessel formation, 

Nature, 376, 70, 1995. 

Scarborough, R. M. and Gretler, D. D., Platelet glycoprotein IIb-IIIa antagonists as prototypical integrin 

blockers: Novel parenteral and potential oral antithrombotic agents, J. Med. Chem., 43, 3453, 2000. 

Schiffelers, R. M. et al., Anti-tumor efficacy of tumor vasculature-targeted liposomal doxorubicin, J. Control. 

Release, 91, 115, 2003. 

Schlingemann, R. O. et al., Leukocyte antigen CD34 is expressed by a subset of cultured endothelial cells and 

on endothelial abluminal microprocesses in the tumor stroma, Lab. Investig., 62, 690, 1990. 

Seon, B. K. et al., Long-lasting complete inhibition of human solid tumors in SCID mice by targeting 

endothelial cells of tumor vasculature with antihuman endoglin immunotoxin, Clin. Cancer Res., 

3,1031,1997. 

Shimaoka, M., Takagi, J., and Springer, T. A., Conformational regulation of integrin structure and function, 

Annu. Rev. Biophys. Biomol. Struct., 31, 485, 2002. 

Springer, T. A., Traffic signals for lymphocyte recirculation and leukocyte emigration: The multi-step paradigm, 

Cell, 76, 301, 1994. 

Suzuki, S. and Naitoh, Y., Amino acid sequence of a novel integrin 34 subunit and primary expression of the 

mRNA in epithelial cells, EMBO J., 9, 757, 1990. 



Tabata, M. et al., Antiangiogenic radioimmunotherapy of human solid tumors in SCID mice using (125)Ilabeled 

anti-endoglin monoclonal antibodies, Int. J. Cancer, 82, 737, 1999. 

Tamura, R. N. et al., Epithelial integrin a4p6: Complete primary structure of a6 and variants forms of P4, J. Cell 

Biol., 111(1593), 604, 1990. 

Thurston, G. et al., Cationic liposomes target angiogenic endothelial cells in tumors and chronic inflammation 

in mice, J. Clin. Invest., 101,1401, 1998. 

Varner, J. A. and Cheresh, D. A., Tumor angiogenesis and the role of vascular cell integrin avb3, Important 

Adv. Oncol., 69, 1996. 

Vyas, S. P. and Khar, R. K., Drug Delivery to Tumor, in Targeted and Controlled Drug Delivery, Novel 

Carrier Systems, CBS Publishers, New Delhi, 2001. 

Xiong, X. B. et al., Enhanced intracellular uptake of sterically stabilized liposomal Doxorubicin in vitro 

resulting in improved antitumor activity in vivo, Pharm. Res., 22, 933, 2005. 

Yamamoto, S. et al., Antimetastatic effects of synthetic peptides containing the core sequence of the type III 

connecting segment domain (IIICS) of fibronectin, Anticancer Drugs, 5, 4424, 1994. 

Yusuf-Makagiansar, H. et al., Inhibition of LFA-1/ICAM-1 and VLA-4/VCAM-1 as a therapeutic approach to 

inflammation and autoimmune diseases, Med. Res. Rev., 22, 146, 2002. 

Zardi, L. et al., Elution of fibronectin proteolytic fragments from a hydroxyapatite chromatography column, 

Eur. J. BioChem., 146, 5711, 1985. 


33 

CONTENTS 

Folate Receptor-Targeted 

Liposomes for Cancer Therapy 

Xiaobin B. Zhao, Natarajan Muthusamy, 

John C. Byrd, and Robert J. Lee 

33.1 Introduction ....................................................................................................................... 663 

33.2 The Folate Receptor as a Cancer-Specific Cellular Marker............................................. 663 

33.3 Targeting FR for Cancer Therapy..................................................................................... 664 

33.4 Liposomes as Drug Carriers.............................................................................................. 665 

33.5 Formulation Strategies for FR-Targeted Liposomes ........................................................ 667 

33.6 In Vitro Drug Delivery Using FR-Targeted Liposomes .................................................. 668 

33.7 Pharmacokinetic and Biodistribution Properties of FR-Targeted Liposomes ................. 670 

33.8 In Vivo Therapeutic Efficacy of FR-Targeted Liposomes ............................................... 671 

33.9 Conclusions and Future Directions ................................................................................... 672 


References..................................................................................................................................... 673 


Targeted drug delivery is a promising strategy to improve both the efficacy and safety of treatment. 

This is especially attractive for cancer therapy as most anti-cancer drugs have only marginal 

therapeutic index. Various cellular surface receptors and antigens have been targeted through 

different strategies, among which folate receptor (FR) has been one of the targets extensively 

investigated. FR is selectively amplified on human malignant cells and can take up folate and its 

analogs through a receptor-mediated endocytosis process. 1,2 FR-targeted liposomes can be utilized 

to deliver therapeutic agents to FR positive cells. 3 These liposomes have been evaluated for the 

targeted delivery of a wide variety of agents. This review will highlight recent developments on 

FR-targeted liposomes with emphasis on the application of this targeted nanoscale delivery system 

in cancer therapy. Potential applications of the FR-targeted liposomes in the clinics will also 

be discussed. 

33.2 THE FOLATE RECEPTOR AS A CANCER-SPECIFIC CELLULAR MARKER 

FR, also known as folate-binding proteins (FBP), is an N-glycosylated protein with high binding 

affinity to folate. FR has three isoforms, a, b, and g/g 0 . FR-a and FR-b are glycosyl-phosphatidylinositol 

(GPI)-anchored membrane bound proteins, 4,5 whereas FR-g/g 0 are constitutively 

secreted. 1,6,7 The g 0 isoform is a truncated form of FR-g. 8 These isoforms display different patterns 


663

664 

of tissue specificity, and they also present differential ligand stereospecificities. 1 The two 

membrane receptor subtypes, a and b, share high amino acid sequence identity (70%) but are 

distinguishable by differential affinities for folic acid and stereoisoforms of reduced folates (affinities 

for folic acid: FR-a K dw0.1 nM, 9 FR- b K dw1 nM, 10 and FR- g K dw0.4 nM 6 ). 

Expression of FR is both tissue-specific and differentiation dependent. The expression of FR 

has been identified by immuno-histochemical staining, reverse transcription-polymerase chain 

reaction (RT–PCR), Western blotting, and 3 H-folic acid binding both in normal and malignant 

tissues. 11,12 Functional FR expression is absent in most normal tissues with the exception of the 

luminal surface of certain epithelial cells 2 where it has limited accessibility to the blood stream. On 

the other hand, FR-a is consistently expressed in several carcinomas, 12 especially in non-mucinous 

ovarian carcinomas, uterine carcinomas, testicular choriocarcinomas, ependymomas, and pleural 

mesotheliomas, and it is less frequently expressed in breast, colon, and renal cell carcinomas. 2,13 

Correlation of FR-a expression level and histologic grade has also been shown in ovarian and 

breast cancers, suggesting FR-a as a possible malignant cellular marker for cancers. 14 Methods for 

upregulating FR-a expression using anti-estrogens and gluococorticoid agonists have recently been 

published. 15,16 The effect of these agents is further enhanced by histone deacetylase (HDAC) 

inhibitors. 16 FR-b is a differentiation marker in the myelomonocytic lineage during neutrophil 

maturation, 17 and it is amplified in activated monocytes and macrophages. 18 However, FR-b in 

neutrophils is unable to bind folate as a result of aberrant post-translational modifications. 19 FR-b 

is also expressed in a functional form in chronic myelogenous leukemia (CML), in 70% of acute 

myelogenous leukemias (AML), 7,17,20,21 and in activated macrophages associated with rheumatoid 

arthritis. 22 FR-b expression is regulated by retinoid receptors and can be upregulated by all-trans 

retinoic acid (ATRA). 23 The effect of ATRA is further enhanced by HDAC inhibitors. The selective 

amplification of FR-a and FR-b expression in human cancers suggests possible roles for FRs 

as cancer specific cellular markers and their potential utility as targets for drug and gene delivery 

to these diseases. FR-g/g 0 are reported to be specific to hematopoietic tissues but are expressed 

at very low levels. 7 FR-g/g 0 are in secreted form, having limited application for targeted drug 

delivery but may serve as a serum marker to monitor certain hematological malignancies. 7 The 

additional possibility of selective FR upregulation in the target cells might further enhance the 

targeting strategy. 

33.3 TARGETING FR FOR CANCER THERAPY 


The selective amplification of FR expression in both human solid tumors and leukemia suggests its 

utility as a potentially valuable target for drug and gene delivery. Both monoclonal antibodies (MAbs) 

against FR and folate itself have been evaluated as moieties for targeting to the FR. 

Two MAbs, MOv18, and MOv19, have been produced against FR-a on ovarian cancers. Both 

aremurineMAbsoftheIgG1 class raised against a poorly differentiated ovarian carcinoma, 

recognizing two distinct epitopes. 24 Clinical studies on radioimmunoscintigraphy using 

131 I-MOv18 have been carried out in ovarian cancer patients and demonstrated targeting of 

ovarian cancer. 25 In addition, a particle-emitter 211 At conjugated to MOv18 has been found to 

prolong survival in a murine peritoneal tumor model developed by intraperitoneal inoculation of 

OVCAR-3 cells. 

Folate itself has, in fact, been the focus of recent development of FR targeting ligand. This is 

possibly because folate is a low molecular weight (MWZ441) ligand and has high affinity to FR, 

the conjugation chemistry of folate is defined and relatively easy to carry out, folate is a compound 

of unlimited availability, and folate lacks immunogenicity in contrast to murine MAb. The conjugation 

of folate via its g-carboxyl has been shown to preserve its binding affinity to the FRs. Recent 

studies also found that the glutamate residue of folic acid is not critical for FR recognition, and only 

pteroic acid is essential for FR binding. 61 Similar to the uptake process of the vitamin folate, FRs 


Folate Receptor-Targeted Liposomes for Cancer Therapy 665 

mediate the cellular internalization of folate conjugates via receptor-mediated endocytosis; 19,20 

therefore, this highly specific event can be utilized to promote uptake of folate-therapeutic agent 

conjugates in FR positive cancer cells. 

Early proof-of-concept in vivo studies were conducted using either radio-labeled or fluorochrome-labeled 

folic acid. The selectivity has also been documented in tumor-bearing mice using 

radiolabeled folate conjugates such as 67 Ga-deferoxamine-folate, 99m Tc-hydrazinonicotinic acid 

(HYNIC)-folate, and 111 In-diethylenetriamine pentaacetic acid (DTPA)-folate as potential folatebased 

radiopharmaceuticals. 26,27 Among these, 111 In-DTPA-folate has been further evaluated in 

human clinical trials. Concentration of radioactivity in abdominal masses was found in women 

suspected of ovarian cancer, and only the kidneys and livers in some patients displayed significant 

retention of 111 In-DTPA-folate. This reagent was developed by Endocyte Inc., an Indiana-based 

biopharmaceutical company, to be a folate-targeted radiopharmaceutical imaging agent 

(FolateScane) and is currently in phase II clinical trial designed to evaluate if FolateScane can 

be used to detect and make a disease or pathological assessment of ovarian cancer. 

A wide range of therapeutic agents have also been evaluated for enhanced cancer cell selective 

delivery. These include chemotherapeutic agents, 28 radiopharmaceuticals, 29 prodrug-converting 

enzymes, 30 protein toxins, 31 T-cell specific antibodies, 32 haptens, 30 gene transfer vectors, 33,34 

antisense oligodeoxyribonucleotides (ODNs), 35 polymeric drug carriers, 36 and liposomes 

19,37,38,44,49,51,56–60 carrying a variety of therapeutic agents. These conjugates potentially 

have broad applications in numerous imaging and therapeutic modalities. 

Based on the action site of therapeutic agents, application of FR-targeted delivery can be 

categorized into two types. For hydrophilic compounds with limited permeability through cell 

membrane, FR offers an efficient drug uptake pathway that can both specify and facilitate the 

accessibility of therapeutic active agents to cancer cells. Folate-conjugates enter FR positive 

cells through receptor-mediated endocytosis that is different from passive permeation process. 

Delivery to cancer cells through a process different from traditional drug transportation has the 

potential to make the drug more accessible to cancer cells. On the other hand, for drugs that are 

ready to take effect on the surface of cells such as immunomodulator and prodrug-activating 

enzymes, FR may also act as a concentrator for these types of drugs. FRs are generally over 

expressed on the cellular surface of various cancer types, and upon binding of FR drug conjugates, 

they may enrich these therapeutic agents on the cell surface. Therefore, application of FR-targeted 

delivery can be exploited for variety of compounds (Figure 33.1). 

Taken together, FR targeting has been investigated as a valuable approach to enhance the 

delivery of variety of agents. Therefore, the interest in exploiting this receptor for targeting application 

is still increasing as demonstrated by the active researches conducted in both academia and 

industry. Among these, FR-targeted liposome, first designed and investigated by Lee and Low, 37,38 

has been attractive because of the advantages it possesses as a useful delivery system both for 

hydrophilic and hydrophobic agents. 

33.4 LIPOSOMES AS DRUG CARRIERS 

Liposomes are spherical, self-closed structures composed of lipophilic bilayers with entrapped 

hydrophilic core. The development of small size, stable, long-circulating, targeted liposomes has 

led to an exciting era in the pharmaceutical application of nanotechnology. Giving the unique 

structure of liposomes compared to other drug delivery systems, both lipophilic and lipophobic 

therapeutic agents can be delivered by liposomes. The biocompatibility of liposomes, if used with 

proper lipid composition, also makes liposomes attractive as safe and effective drug carriers. A 

variety of therapeutic agents have been incorporated into liposomes. Several have reached clinical 

use. These include liposomal doxorubicin (Doxile), daunorubicin (Daunoxomee), amphotericin B 

(Amphotece, Ambisomee, Abelcete), cytarabine (Depocytee), and verteporfin (Visudynee). 


666 

OH 

N 

H N 

2N N 

N 

Folic acid 

NH 

α 

O 

COOH 

NH 

COOH γ 

FIGURE 33.1 Folate-conjugates for FR-targeted delivery. 

Numerous liposomal formulations are in clinical trial, including vincristine, all-trans retinoid acid, 

topotecan, and cationic liposome-based therapeutic gene transfer vectors. Many more are in preclinical 

evaluation. Besides potential use in systemic gene delivery, cationic liposomes are routinely 

used as transfection reagents for plasmid DNA and oligonucleotides in the laboratory. The use of 

liposomes can be deliberately engineered to possess unique properties such as long systemic 

circulation time, pH sensitivity, temperature sensitivity, and target cell specificity. These are 

achieved by selecting the appropriate lipid composition and surface modification of the liposomes, 

expanding the application of liposomes for drug delivery. 

Liposomal delivery of anti-cancer drugs has been shown to greatly extend their systemic 

circulation time, reduce toxicity by lowering plasma free drug concentration, and facilitate preferential 

localization of drugs in solid tumors based on increased endothelial permeability and reduced 

lymphatic drainage or enhanced permeability and retention (EPR) effect. 39 Clearance of drugs in a 

liposomal formulation is mediated by phagocytic cells of the reticuloendothelial system (RES), 

primarily located in the liver and spleen. Drugs can be loaded during liposome preparation through 

methods such as passive entrapment. Alternatively, they can also be loaded after liposome preparation 

through a process called remote loading, utilizing the ion gradient across the membrane as 

in the case of doxorubicin. 31 Compared to free drugs, liposomal drugs are not subjected to rapid 

renal clearance. Instead, they are cleared by RES, exhibiting prolonged systemic circulation time. 40 

RES clearance of liposomes can be reduced by surface coating with polyethyleneglycol (PEG). 

Passive targeting effect has also been shown in liposomal drug delivery to solid tumors where 

liposomes preferentially localize in these tumor tissues as a result of the EPR effect. 41 

Liposomes can be targeted to specific cell populations via incorporation of a desired targeting 

moiety such as a chemical conjugated ligand, antibody, or antibody fragment directed against the 

surface molecules. Targeted delivery of these liposomes can greatly improve their selectivity of 

cancer cells and facilitate their cellular uptake and intracellular drug release. Examples of such 


E 


Radiopharmaceuticals 

Chemotherapeutics 

DNA and antisense ODN 

Prodrug converting enzymes 

Antibodies 

BNCT agents 

Polymeric drug carriers 

Liposomes


targeted liposomes are folate-conjugated liposomes targeting to acute myelogenous leukemia, 

CD19-targeted immunoliposomes for non-Hodgkin’s lymphoma (NHL) therapy, 42 and 

anti-HER2 immunoliposomal doxorubicin targeting to HER2-overexpressing breast cancer 

cells. 43 A variety of techniques have been described to incorporate targeting moieties to 

liposomes. First, ligands such as folate can beconjugatedtoalipidcomponent suchas 

poly(ethylene glycol)-conjugated distearoyl phosphatidylethanolamine (PEG-DSPE) and PEG- 

Chol. This amphophilic lipid component can be constructed into the lipid bilayer during the 

formation of membrane leaving the ligand on the surface of the particles. 38,59 In the second 

method, a functionalized lipid anchor is included first to the outside of the liposomes after the 

formation of the bilayer membrane. Therefore, the exposed crosslinker can react with activated 

targeting moieties such as thiolated proteins (antibody, scFv, transferrin, etc.). 42,43 A useful thiolation 

reagent is 2-iminothiolane that is also known as Traut’s reagent. There are two types of 

immunoliposomes based on the antibody coupling strategy. Type one involves direct attachment 

of the antibodies to the lipids, and type two involves linking the antibodies to the terminal ends of 

reactive PEG derivates. Most of these coupling techniques lead to a random antibody orientation at 

the liposome surface. A significant contribution to the development of the type two immunoliposome, 

recently adopted for formulation of HER2-targeted immunoliposomes, is an antibody 

coupling method called post-insertion method. 43 This method involves formation of micelles of 

lipid-derivatized antibodies first, followed by transferring the antibody-coupled PEG-lipid micelles 

to the preformed liposomes, converting the conventional liposomes to immunoliposome through a 

simple one-step incubation method. Maleimide-terminated PEG–DSPE is a coupling lipid that has 

been validated to successfully convert commercially available liposomal doxorubicin (Doxil/ 

Caelyx) to targeted sterically stabilized liposomal doxorubicin. The last post-insertion method 

seems highly promising for future clinical development. 

33.5 FORMULATION STRATEGIES FOR FR-TARGETED LIPOSOMES 

FR-targeted liposomes can be prepared by including a small fraction (e.g., 0.1 mole%) of a lipophilic 

folate derivative into the lipid composition. 38 Both folate-PEG (M.W. 3350)-distearoyl 

phosphatidylethanolamine (folate-PEG–DSPE) 38 and folate-PEG-cholesterol (folate-PEG-chol) 

have been synthesized and shown to be effective in targeting liposomes to the FR. The synthesis 

of folate-PEG–DSPE and folate-PEG-Chol can be realized through two stages. First, an intermediate 

folate-PEG-amine is synthesized by reacting NHS-activated folate and reacted with PEG-bis 

amine and further purifized using a Sepharose column. This folate-PEG-NH 2 is then ready to react 

with N-succinyl-DSPE to form folate-PEG–DSPE 38 or with cholestreryl chloroformate to form 

folate-PEG-chol. 59 Another synthetic method has also been described by Gabizon that is to couple 

FA to readily attainable H2N-PEG–DSPE with the mediation of carbodiimide. 44 These two synthesized 

lipophilic folate-derivatives have both been shown to be good components to be 

incorporated to liosomes. Compared to the synthesis of folate-PEG–DSPE, the synthesis of 

folate-PEG-chol is relatively simple and is less costly. With regard to stability, cholesterol also 

provides better chemical stability, and it is also known to be able to increase the structural integrity 

of lipid membranes through tight packing. The molecule of cholesterol is neutrally charged, 

different from the negatively charged DSPE molecule, and can possibly provide better targeting 

effect when used as a lipophilic anchor to incorporate the folate molecule. For these reasons, this lab 

has been routinely using folate-PEG-chol for the formulation of FR-targeted liposomes. 

Preparation of folate receptor (FR) targeted liposomes is relatively straightforward since the 

targeting lipid component can simply be included into the lipid composition during the formation of 

liposomes. A small percentage (0.1–0.5%, molar ratio) of folate-PEG–DSPE or folate-PEG-chol in 

the lipid can yield targeting effect to FR. Relatively long linker between folate and the lipid anchor 

(e.g., PEG 3350) was found to be necessary for FR binding of the liposomes. 37 FR targeting was 

compatible with incorporation of 3 mole% of PEG–DSPE that is required for prolonging the in vivo 


668 


circulation time of the liposomes. Lab scale production of these targeted liposomes can be 

performed by thin-layer hydration followed by polycarbonate membrane extrusion. For scaling 

up, alternative methods such as liquid phase emulsification, high pressure homogenization, and 

tangential flow filtration can be used. Recently, this lab has successfully validated a processing 

method combining homogenization, ultrafiltation, and remote loading for FR-targeted liposomal 

doxorubicin at clinical relevant scale quantity. Using this method, three consecutive batches of 

liposomes with mean particle size around 100 nm curve produced. Another alternative method for 

folate-PEG-Chol incorporation has not yet been investigated, using post-insertion method similar to 

HER2-targeted immunoliposome production by incubating micelles of folate-PEG-chol/PEG-chol 

to pre-formed liposomal drugs. The latter method avoided the ligand stability concern during the 

liposome production and has the potential to be used for future Current Good Manufacturing 

Practices (cGMP) grade production of FR-targeted liposomes. Both hydrophilic drugs and hydrophobic 

drugs can be potentially loaded to the targeted particles through remote loading 

(doxorubicin, daunorubicin, vincristine, and topetecan) or passive entrapment (paclitaxel and 

ATRA); therefore, this FR-targeted liposomal drug delivery represents a valuable approach for 

FR-positive cancer therapy (Figure 33.2). 

FR-targeted cationic liposome formulations have also been studied for gene delivery. Several 

cationic liposome formulations that incorporate a lipopholic folate derivative as a targeting entity 

have been reported. FR-targeted lipoplexes constructed using a cationic lipid RPR209120 was 

shown to efficiently transfer gene. 45 Another study using 2% DPPE–PEG 3340-folate and a cationic 

dithiol-detergent (dimerized tetradecyl-ornithinly-cysteine) demonstrated efficient FR-dependent 

cellular uptake and transfection. Therapeutic gene p53 was also delivered by Xu et al., using 

cationic liposomes conjugated to folate, and this was shown to enhance the anti-tumor efficacy 

of conventional chemotherapy. 46 FR-targeted cationic liposomes can further form complex with 

DNA-condensing polymer to form FR-targeted lipopolyplexes, and this was demonstrated by 

Reddy et al. 54 to have superior transfection activity in FRCM109 murine lung carcinoma cells as 

well as in L1210A murine lymphocytic leukemia cells. Lee et al. developed FR-targeted type II 

lipopolyplexes, and these FR-targeted vectors exhibited improved gene transfer activity even in the 

presence of serum. 47,48 Leamon et al. also investigated the use of FR-targeted liposome constructed 

with lipid-PEG-pteroate for effective delivery of antisense oligonucleotides into FR-bearing cells 

in vitro. 61 FR-targeted liposomal drug and gene delivery has also been the subject of another review 

article. 35 

33.6 IN VITRO DRUG DELIVERY USING FR-TARGETED LIPOSOMES 

Cellular uptake of FR-targeted liposomes has been characterized using KB cells, a FR-a (C) 

human oral carcinoma cell line. 37 Binding of the liposomes has a relatively slow kinetics, occurring 

over several hours, presumably a result of their size. The specific binding of these liposomes to KB 

cells was saturable and could be blocked by excess free folate in the binding media. 37 Interestingly, 

the apparent affinity of the folate-conjugated liposomes to KB cells was much higher than that of 

free folate. 37 This is due to multivalent binding of the folate ligand on the liposomes with the FRs 

on the cellular surface. 37 The folate-liposomes are internalized into a low pH compartment via 

receptor-mediated endocytosis, and the encapsulated drug molecules are released into the cytoplasm 

to induce cytotoxicity 47 as illustrated in Figure 33.3. 

Drug delivery properties of FR-targeted liposomes have been studied in vitro using liposomes 

loaded with chemotherapeutic agents such as doxorubicin, 38 daunorubicin, 49 and cisplatin. 44 Lee 

et al. first reported the in vitro effect of doxorubin and showed these targeted liposomes showed 

w86-fold greater cytotoxicity in KB cells compared to non-targeted control liposomes. 38 The 

enhancement in cytotoxicity was correlated with the increase in doxorubicin uptake and could 

be blocked by excess free folate. 38 Furthermore, a study by Goren et al. showed that the delivery of 



H 2 N 

N 

H 2 N 

OH 

N 

N 

N 

N 

OH 

N 

NH 

N 

N 

NH 

Folate 

Folate receptor 

HOOC 

NH 

O 

O 

Folate-PEG-cholesterol 

COOH PEG 

O 

NH NH 

NH 

O 

O 

NH 

O 

Folate-PEG-DSPE 

PEG 2000 

PEG O 

NH 

NH 

O 

O P O 

Hydrophobic reagents 

Hydrophilic reagents 

FR-targeted liposome 

Cancer cell membrane 

FIGURE 33.2 FR-targeted liposome. The ligand folate is conjugated to a lipophilic anchor (PEG3350-DSPE or 

PEG3350-chol) and incorporated to the lipid bilayer. PEG2000–DSPE is also included in the construction of the 

liposome membrane, prolonging the systemic circulation of these particles by avoiding the RES uptake. 

Hydrophilic anti-cancer reagents can be loaded to the aqueous core through active loading or passive entrapment, 

and hydrophobic reagents can be loaded to the lipid bilayer through formation of electrostatic complex 

or passive entrapment. The structures of two synthetic targeting lipids (F-PEG–DSPE and F-PEG-Chol) were 

also shown. 

doxorubicin via FR-targeted liposomes bypassed Pgp-dependent drug efflux in drug resistant 

FR-a(C) M109 murine lung carcinoma cells in vitro and in an adoptive assay in mice. 50 These 

data suggest that this delivery modality might overcome drug resistance in these tumor cells. 50 

More recently, folate-liposomal doxorubicin was studied in FR-b(C) KG1 human AML cells and 

similarly showed enhanced cytotoxicity relative to non-targeted control liposomes, 19 and the effect 


O- 

O 

O 

O 

O 

O

670 

FIGURE 33.3 FR-mediated endocytosis of folate-liposomes. 

was enhanced by FR-b upregulation using ATRA. 19 Therefore, both FR-a and FR-b are potential 

targets for folate-conjugated liposomes. 

In addition to chemotherapeutics, a variety of agents have been loaded into FR-targeted 

liposomes and evaluated in vitro. These include fluorescent 37 probes, BNCT agents, 51 photosensitizers, 

52 antisene oligodeoxyribonucleotides against the EGF receptor, 53 and plasmid DNA 

carrying luciferase reporter gene. 47,54 Similar formulations of FR-targeted lipid nanoparticles 

containing paclitaxel have also been recently reported. 55 In each of these studies, the FR-targeted 

formulationwasshowntobemuchmoreefficient in cellular uptake and biological activity 

compared to the non-targeted control formulation. 

33.7 PHARMACOKINETIC AND BIODISTRIBUTION PROPERTIES 

OF FR-TARGETED LIPOSOMES 


The in vivo clearance rate for folate-conjugated liposomal doxorubicin was faster than non-targeted 

control liposomes in mice engrafted with FR-overexpressing M109 or human KB carcinoma or 

mouse J6453 lymphoma. 56 These results also suggested a possible role of FR in the differential 

rates of clearance. 56 Plasma clearance kinetics of FR-targeted liposomes radiolabled with 67 Ga 

were studied in rats. These liposomes showed enhanced plasma clearance rates and higher uptakes 

in the liver and spleen compared to non-targeted control liposomes. 56 This effect was even more 

pronounced in rats that were placed on a folate-free diet, suggesting that FR expression in the RES 

organs might be responsible for the increase in the rate of liposome clearance. 56 One possible explanation 

is that FR-targeted liposomes, by virtue of their multivalency, can interact with the FR-b that 

is present at a relatively low level in the phagocytic cells. 

In the biodistribution studies in murine tumor models, FR-targeted liposomes did not show 

enhanced accumulation in tumors relative to non-targeted liposomes. 59 This might be due to the 

passive accumulation of the liposomes via the EPR effect was the predominant factor in tumor 

localization of both targeted and non-targeted liposomes because endothelial extravasation was 

the rate-limiting step in tumor localization. In addition, FR-mediated enhancement in uptake of 

the targeted liposomes might be offset by the more rapid plasma clearance of these liposomes as 

discussed above. 60 Moreover, in solid tumors, FRs on tumor cell surface are not directly exposed 

to the circulation and not fully accessible by circulating FR-targeted liposomes. Therefore, extravasation, 

instead of specific binding, is the rate-limiting factor of liposome distribution in 

solid tumors. 



An important observation in vivo is that macrophages also have high uptake of FRtargeted 

liposomes. In a study done in an ovarian carcinoma ascites mouse model, results 

showed that macrophages took up FR-targeted liposomes efficiently in the ascites fluid even in 

the presence of FR positive ascetic tumor cells, 57 although other reports did show increased 

FR-targeted liposome uptake in FR(C) ascitic tumor cells. 56 The relative uptake of folateliposomes 

by these two cell populations was likely determined by the relative FR expression 

levels. 56 

33.8 IN VIVO THERAPEUTIC EFFICACY OF FR-TARGETED LIPOSOMES 

The therapeutic efficacy of folate-liposomal doxorubicin has been evaluated in several murine 

tumor models. In a KB cell xenograft model of athymic mice on a folate-free diet, folate-liposomal 

doxorubicin given i.p. at 10 mg/kg!6 injections showed significantly greater tumor inhibition 

compared to non-targeted liposomes. 58 However, in another study using a M109R-HiFR syngraft 

model in BALB/c mice on a regular folate-containing diet, folate-liposomal doxorubicin given 

intravenously at 8 mg/kg!4 injections showed a similar therapeutic efficacy to non-targeted 

control liposomes. 3 The differences in these studies might be partly due to the different routes of 

administration, animal diet, and dosages applied. 3 A challenge in demonstrating a therapeutic 

benefit of folate-liposomal doxorubicin is the excellent anti-tumor activity of PEGylated liposomal 

doxorubicin used as a control that has optimal pharmacokinetic properties and EPR effect and are 

much more effective than the free drug. 3 Further studies in additional models are likely needed to 

confirm possible therapeutic benefits of FR-targeted liposomal doxorubicin. 

A possible mechanism for increased therapeutic efficacy for the folate-liposomal doxorubicin in 

solid tumor is altered intratumoral drug distribution. These liposomes might be more efficiently 

internalized by FR(C) tumor cells and/or tumor infiltrating macrophages that are FR(C). 57 This, in 

turn, results in an increase in drug release within the tumor. In contrast, non-targeted liposomal 

doxorubicin may remain extracellular and be distributed in the interstitial space and releases drug 

more slowly. This may result in a bioavailable (free) drug concentration that falls below the 

minimum effective concentration (MEC) for inducing tumor cell apoptosis. Therefore, cellular 

internalization and associated drug release may be a critical mechanism of liposomal drug action 

in solid tumors and a rate-limiting factor in overall drug delivery pathway. The relative importance 

of FR expression on tumor cells and tumor infiltrating macrophages and their role in uptake of 

folate-coated liposomes warrant further investigation. 

Leukemia is a better disease target as compared to solid tumors for this targeted delivery 

strategy. This potentially is because malignant cells are circulating in leukemia patients and are 

fully accessible by targeted liposomes. In contrast, these liposomes need to overcome several 

barriers in solid tumors before reaching cancer cells. Several studies have already reported the 

anti-tumor efficacy of FR-targeted liposomal anthracyclines in leukemia. In both L1210JF murine 

lymphocytic leukemia ascites models and a KG-1 human acute myelogenous leukemia xenograft 

model, FR-targeted liposomal doxorubicin or daunorubicin, when administered i.p, 19 showed 

improvement in survival of treated mice compared to non-targeted control liposomes or the free 

drug in the L1210JF model. 19 In both mice models, folate-liposomal doxorubicin given i.p. was 

also more effective in prolonging animal survival. 19 The relatively greater accessibility of 

leukemia cells in ascites tumors might have contributed to the superior efficacy of the 

targeted liposomes. 

Interestingly, pretreatment of mice with ATRA, an agent known to upregulate FR-b expression 

in KG-1 cells in vitro, 19 before the injection of liposomes was also shown to enhance the 

efficacy. An increase in long-term survival in these mice as a result of FR-targeted liposomal 

doxorubicin was shown from 12.5% to 60%. Therefore, administrating ATRA to upregulate 


672 

FR-b expression seems to be a potentially promising therapeutic strategy to be combined with the 

FR-targeted therapy. 

33.9 CONCLUSIONS AND FUTURE DIRECTIONS 


Targeted drug delivery is a promising approach to selectively deliver neoplastic agents to cancer 

cells and to improve the therapeutic index of chemotherapy. FR outstands among various tumor 

markers because of the specificity and relatively high level of endocytosis, representing a very 

attractive target for developing targeted delivery for future cancer therapy. FR-targeted liposomal 

delivery is a promising strategy for treatment of FR(C) tumors and leukemias. Although a variety 

of agents have been delivered to FR(C) tumor cell in vitro using these liposomes, relatively few 

reports address the in vivo properties of these liposomes. However, at least some studies indicated a 

therapeutic advantage of folate-liposomal doxorubicin over the non-targeted control. 19,49,58 

Leukemia, that has greater accessibility from the systemic circulation, may present an even 

better therapeutic target for these liposomes. The concept of sensitizing AML cells with 

FR-b-upregulating agents such as ATRA may bring further rationale to take FR-targeted liposomal 

drug delivery to clinical trials. 

Several critical factors to be considered for designing in vivo application of FR-targeted liposomes 

for cancer therapy include FR expression level on tumor cells relative to normal cells, the 

accessibility of liposomes to these receptors, the intra-tumor diffusion rate of vectors, and the 

presence of endogenous folate in systemic circulation. The first obstacle of insufficient levels of 

FR expression on tumor cells in patients could be possibly overcome by inducing FR expression 

upregulation by co-administration of anti-estrogen or retinoid receptor ligands. To address the 

second concern of accessibility, optimization of formulation parameter is necessary to reduce the 

particle size, reduce the plasma protein binding, and avoid the RES and other normal tissue uptake. 

For solid tumor delivery, priming the vasculature in the tumor to increase the distribution of 

targeted particles is a possible approach. The concern that the presence of endogenous folate in 

systemic circulation can block FR binding is not a critical one. It is believed that this should not 

constitute a barrier as the endogenous form of folate has lower affinity for the FR compared to folic 

acid and should not significantly impede the binding of multivalent folate-coated liposomes. 

Clinical success in targeted drug delivery area has so far been limited by the lack of a suitable 

cellular target and/or a high degree of difficulty in the production of clinical quantities of targeted 

drug carriers. FR-targeted liposomes represent an attractive candidate for clinical development with 

potential application in the treatment of AML and other FR-over expressing solid tumors. Areas of 

investigation for future studies include the possible role of receptor upregulation in promoting 

targeted therapy, both for FR-a and FR-b positive tumors; and the cellular population responsible 

for enhancement in therapeutic response in solid tumors, i.e., relative roles of tumor infiltrating 

macrophages and the tumor cells themselves. 57 Because activated macrophages have increased FRb 

expression, 22 it is possible that folate-conjugates can effectively target tumor infiltrating macrophages 

in non-FR expressing tumors, extending the FR-targeting strategy to receptor 

negative tumors. 

The area of FR targeting remains an area of great excitement, suggested by large number of 

recent publications and reports from industrial development. So far, no clinical study targeting FR 

for therapeutic purposes has been carried out. Given the advantages of the FR-targeting strategy, 

further preclinical studies are urgently needed to define mechanisms of in vivo targeting and the 

potential for clinical efficacy. FR-targeted liposomal doxorubicin represents an compelling candidate 

for clinical development with potential application in the treatment of AML. Furthermore, 

future application to other tumors with high frequency of FR overexpression such as ovarian and 

lung carcinomas warrants investigation. 




This work was supported, in part, by NCI grant ROI CA095673 to RJL, P01CA95426 to JCB, NSF 

grants EEC-0425626 to RJL, and Leukemia and Lymphoma Society grant 6113-02 to RJL. 

REFERENCES 

1. Elnakat, H. and Ratnam, M., Distribution, functionality and gene regulation of folate receptor 

isoforms: Implications in targeted therapy, Advanced Drug Delivery Reviews, 56(8), 1067–1084, 

2004. 

2. Weitman, S. D. et al., Distribution of the folate receptor GP38 in normal and malignant cell lines and 

tissues, Cancer Research, 52(12), 3396–3401, 1992. 

3. Gabizon, A. et al., Tumor cell targeting of liposome-entrapped drugs with phospholipid-anchored 

folic acid-PEG conjugates, Advanced Drug Delivery Reviews, 56(8), 1177–1192, 2004. 

4. Lacey, S. W. et al., Complementary DNA for the folate-binding protein correctly predicts anchoring 

to the membrane by glycosyl-phosphatidylinositol, The Journal of Clinical Investigation, 84(2), 

715–720, 1989. 

5. Ratnam, M. et al., Homologous membrane folate-binding proteins in human placenta: Cloning and 

sequence of a cDNA, Biochemistry, 28(20), 8249–8254, 1989. 

6. Shen, F. et al., Folate receptor type gamma is primarily a secretary protein due to lack of an efficient 

signal for glycosyl-phosphatidylinositol modification: Protein characterization and cell type specificity, 

Biochemistry, 34(16), 5660–5665, 1995. 

7. Shen, F. et al., Identification of a novel folate receptor, a truncated receptor, and receptor type beta in 

hematopoietic cells: cDNA cloning, expression, immunoreactivity, and tissue specificity, Biochemistry, 

33(5), 1209–1215, 1994. 

8. Wang, H., Ross, J. F., and Ratnam, M., Structure and regulation of a polymorphic gene encoding 

folate receptor type gamma/gamma 0 , Nucleic Acids Research, 26(9), 2132–2142, 1998. 

9. Kamen, B. A. and Caston, J. D., Properties of a folate-binding protein (FBP) isolated from porcine 

kidney, Biochemical Pharmacology, 35(14), 2323–2329, 1986. 

10. da Costa, M. and Rothenberg, S. P., Purification and characterization of folate-binding proteins from 

rat placenta, Biochimica Et Biophysica Acta, 1292(1), 23–30, 1996. 

11. Ross, J. F., Chaudhuri, P. K., and Ratnam, M., Differential regulation of folate receptor isoforms in 

normal and malignant tissues in vivo and in established cell lines, Physiologic and clinical implications. 

Cancer, 73(9), 2432–2443, 1994. 

12. Weitman, S. D. et al., Cellular localization of the folate receptor: Potential role in drug toxicity and 

folate homeostasis, Cancer Research, 52(23), 6708–6711, 1992. 

13. Garin-Chesa, P. et al., Trophoblast and ovarian cancer antigen LK26. Sensitivity and specificity in 

immunopathology and molecular identification as a folate-binding protein, American Journal of 

Pathology, 142(2), 557–567, 1993. 

14. Toffoli, G. et al., Overexpression of folate-binding protein in ovarian cancers, International Journal of 

Cancer, 74(2), 193–198, 1997. 

15. Kelley, K. M. M., Rowan, B. G., and Ratnam, M., Modulation of the folate receptor alpha gene by the 

estrogen receptor: Mechanism and implications in tumor targeting, Cancer Research, 63(11), 

2820–2828, 2003. 

16. Tran, T. et al., Enhancement of folate receptor alpha expression in tumor cells through the glucocorticoid 

receptor: A promising means to improved tumor detection and targeting, Cancer Research, 

65(10), 4431–4441, 2005. 

17. Ross, J. F. et al., Folate receptor type beta is a neutrophilic lineage marker and is differentially 

expressed in myeloid leukemia, Cancer, 85(2), 348–357, 1999. 

18. Nakamura, M., Nitration and chlorination of folic acid by peroxynitrite and hypochlorous acid, and the 

selective binding of 10-nitro-folate to folate receptor beta, Biochemical and Biophysical Research 

Communications, 297(5), 1238–1244, 2002. 

19. Pan, X. Q. et al., Strategy for the treatment of acute myelogenous leukemia based on folate receptor 

beta-targeted liposomal doxorubicin combined with receptor induction using all-trans retinoic acid, 

Blood, 100(2), 594–602, 2002. 


674 


20. Sadasivan, E. et al., Purification, properties, and immunological characterization of folate-binding 

proteins from human leukemia cells, Biochimica Et Biophysica Acta, 925(1), 36–47, 1987. 

21. Sadasivan, E. et al., Characterization of multiple forms of folate-binding protein from human 

leukemia cells, Biochimica Et Biophysica Acta, 882(3), 311–321, 1986. 

22. Paulos, C. M. et al., Folate receptor-mediated targeting of therapeutic and imaging agents to activated 

macrophages in rheumatoid arthritis, Advanced Drug Delivery Reviews, 56(8), 1205–1217, 2004. 

23. Wang, H. et al., Differentiation-independent retinoid induction of folate receptor type beta, a potential 

tumor target in myeloid leukemia, Blood, 96(10), 3529–3536, 2000. 

24. Miotti, S. et al., Characterization of human ovarian carcinoma-associated antigens defined by novel 

monoclonal antibodies with tumor-restricted specificity, International Journal of Cancer, 39(3), 

297–303, 1987. 

25. Kalofonos, H. P., Karamouzis, M. V., and Epenetos, A. A., Radioimmunoscintigraphy in patients with 

ovarian cancer, Acta Oncology, 40(5), 549–557, 2001. 

26. Mathias, C. J., et al., Tumor-selective radiopharmaceutical targeting via receptor-mediated endocytosis 

of gallium-67-deferoxamine-folate, Journal of Nuclear Medicine, 37(6), 1003–1008, 1996. 

27. Mathias, C. J., et al., Indium-111-DTPA-folate as a potential folate-receptor-targeted radiopharmaceutical, 

Journal of Nuclear Medicine, 39(9), 1579–1585, 1998. 

28. Leamon, C. P. and Reddy, J. A., Folate-targeted chemotherapy, Advanced Drug Delivery Reviews, 

56(8), 1127–1141, 2004. 

29. Ke, C. Y., Mathias, C. J., and Green, M. A., Folate-receptor-targeted radionuclide imaging agents, 

Advanced Drug Delivery Reviews, 56(8), 1143–1160, 2004. 

30. Lu, J. Y. et al., Folate-targeted enzyme prodrug cancer therapy utilizing penicillin-V amidase and a 

doxorubicin prodrug, Journal of Drug Targeting, 7(1), 43–53, 1999. 

31. Leamon, C. P. and Low, P. S., Selective targeting of malignant cells with cytotoxin-folate conjugates, 

Journal of Drug Targeting, 2(2), 101–112, 1994. 

32. Roy, E. J. et al., Folate-mediated targeting of T cells to tumors, Advanced Drug Delivery Reviews, 

56(8), 1219–1231, 2004. 

33. Mislick, K. A. et al., Transfection of folate-polylysine DNA complexes: Evidence for lysosomal 

delivery, Bioconjugate Chemistry, 6(5), 512–515, 1995. 

34. Gottschalk, S. et al., Folate receptor mediated DNA delivery into tumor cells: Potosomal disruption 

results in enhanced gene expression, Gene Therapy, 1(3), 185–191, 1994. 

35. Zhao, X. B. and Lee, R. J., Tumor-selective targeted delivery of genes and antisense oligodeoxyribonucleotides 

via the folate receptor, Advanced Drug Delivery Reviews, 56(8), 1193–1204, 

2004. 


dendrimers as potential agents for neutron capture therapy, Bioconjugate Chemistry, 14(1), 158–167, 

2003. 

37. Lee, R. J. and Low, P. S., Delivery of liposomes into cultured KB cells via folate receptor-mediated 

endocytosis, Journal of Biological Chemistry, 269(5), 3198–3204, 1994. 

38. Lee, R. J. and Low, P. S., Folate-mediated tumor cell targeting of liposome-entrapped doxorubicin 

in vitro, Biochimica Et Biophysica Acta, 1233(2), 134–144, 1995. 


blood and enhanced uptake by tumors, Proceedings of the National Academy of Sciences of the United 

States of America, 85(18), 6949–6953, 1988. 

40. Abraham, S. A. et al., The liposomal formulation of doxorubicin, Methods in Enzymology, 391, 

71–97, 2005. 

41. Uster, P. S. et al., Insertion of poly(ethylene glycol) derivatized phospholipid into pre-formed liposomes 

results in prolonged in vivo circulation time, FEBS Letters, 386(2,3), 243–246, 1996. 

42. Allen, T. M., Mumbengegwi, D. R., and Charrois, G. J. R., Anti-CD19-targeted liposomal doxorubicin 

improves the therapeutic efficacy in murine B-cell lymphoma and ameliorates the toxicity of 

liposomes with varying drug release rates, Clinical Cancer Research, 11(9), 3567–3573, 2005. 

43. Park, J. W. et al., Anti-HER2 immunoliposomes: Enhanced efficacy attributable to targeted delivery, 

Clinical Cancer Research, 8(4), 1172–1181, 2002. 

44. Gabizon, A., Targeting folate receptor with folate linked to extremities of poly(ethylene glycol)grafted 

liposomes: In vitro studies, Bioconjugate Chemistry, 10(2), 289–298, 1999. 



45. Hofland, H. E. J. et al., Folate-targeted gene transfer in vivo, Molecular Therapy, 5(6), 739–744, 2002. 

46. Xu, L., Pirollo, K. F., and Chang, E. H., Tumor-targeted p53-gene therapy enhances the efficacy of 

conventional chemo/radiotherapy, Journal of Controlled Release: Official Journal of the Controlled 

Release Society, 74(1–3), 115–128, 2001. 

47. Lee, R. J. and Huang, L., Folate-targeted, anionic liposome-entrapped polylysine-condensed DNA for 

tumor cell-specific gene transfer, The Journal of Biological Chemistry, 271(14), 8481–8487, 1996. 

48. Gosselin, M. A., Guo, W., and Lee, R. J., Incorporation of reversibly cross-linked polyplexes into 

LPDII vectors for gene delivery, Bioconjugate Chemistry, 13(5), 1044–1053, 2002. 

49. Pan, X. Q. and Lee, R. J., In vivo antitumor activity of folate receptor-targeted liposomal daunorubicin 

in a murine leukemia model, Anticancer Research, 25(1A), 343–346, 2005. 

50. Goren, D. et al., Nuclear delivery of doxorubicin via folate-targeted liposomes with bypass of multidrug-resistance 

efflux pump, Clinical Cancer Research, 6(5), 1949–1957, 2000. 

51. Pan, X. Q., Wang, H., and Lee, R. J., Boron delivery to a murine lung carcinoma using folate receptortargeted 

liposomes, Anticancer Research, 22(3), 1629–1633, 2002. 

52. Qualls, M. M. and Thompson, D. H., Chloroaluminum phthalocyanine tetrasulfonate delivered via 

acid-labile diplasmenylcholine-folate liposomes: Intracellular localization and synergistic phototoxicity, 

International Journal of Cancer, 93(3), 384–392, 2001. 

53. Wang, S. et al., Delivery of antisense oligodeoxyribonucleotides against the human epidermal growth 

factor receptor into cultured KB cells with liposomes conjugated to folate via polyethylene glycol, 

Proceedings of the National Academy of Sciences of the United States of America, 92(8), 3318–3322, 

1995. 

54. Reddy, J. A. et al., Folate-targeted, cationic liposome-mediated gene transfer into disseminated 

peritoneal tumors, Gene Therapy, 9(22), 1542–1550, 2002. 

55. Stevens, P. J., Sekido, M., and Lee, R. J., A folate receptor-targeted lipid nanoparticle formulation for 

a lipophilic paclitaxel prodrug, Pharmaceutical Research, 21(12), 2153–2157, 2004. 

56. Gabizon, A. et al., In vivo fate of folate-targeted polyethylene-glycol liposomes in tumor-bearing 

mice, Clinical Cancer Research, 9(17), 6551–6559, 2003. 

57. Turk, M. J., Waters, D. J., and Low, P. S., Folate-conjugated liposomes preferentially target macrophages 

associated with ovarian carcinoma, Cancer Letters, 213(2), 165–172, 2004. 

58. Pan, X. Q., Wang, H., and Lee, R. J., Antitumor activity of folate receptor-targeted liposomal doxorubicin 

in a KB oral carcinoma murine xenograft model, Pharmaceutical Research, 20(3), 417–422, 

2003. 

59. Guo, W., Lee, T., Sudimack, J., and Lee, R. J., Receptor targeted delivery of liposomes via folate- 

PEG-chol, J. Liposome Res., 10, 179–195, 2000. 

60. Gabizon, A., Huberty, J., Straubinger, R. M., Price, D. C., and Papahadjopoulos, D., An improved 

method for in vivo tracing and imaging liposome using a 67 Gallium-deferoxamine complex, Journal 

of Liposome Research, 1, 123–135, 1998. 

61. Leamon, C. P., Cooper, S. R., and Hardee, G. E., Folate-liposome-mediated antisense oligodeoxynucleotide 

targeting to cancer cells: Evaluation in vitro and in vivo, Bioconjugate Chemistry, 14(4), 

738–747, 2003. 


34 

CONTENTS 

Nanoscale Drug Delivery 

Vehicles for Solid Tumors: 

A New Paradigm for Localized 

Drug Delivery Using 

Temperature Sensitive 

Liposomes 

David Needham and Ana Ponce 

34.1 Introduction: The Problem and Current Solutions ........................................................... 678 

34.2 The Challenges Facing Nanotherapeutic Intervention: A Brief Review 

of Lipid-Based Nanocarriers ............................................................................................. 682 

34.2.1 Getting the Drug In ............................................................................................. 682 

34.2.2 Getting the Nanocapsules to the Tumor ............................................................. 684 

34.2.3 Mechanisms of Liposomal Drug Delivery ......................................................... 687 

34.3 New Capsular Design for Delivery of Drugs to and Release of Drugs 

in Localized Solid Tumors................................................................................................ 691 

34.3.1 Getting the Drug Out .......................................................................................... 691 

34.3.2 A New Liposome Design.................................................................................... 693 

34.3.3 Nanostructures and Membrane Permeability...................................................... 695 

34.3.4 In Vivo Growth Delay Studies ........................................................................... 701 

34.4 A New Paradigm for Drug Delivery to Tumor Tissues ................................................... 705 

34.4.1 Vascular Shutdown ............................................................................................. 705 

34.4.2 Exploring Other Tumor Models ......................................................................... 707 

34.4.3 Regional Deposition of Drugs Observed Using MRI ........................................ 709 

34.4.4 Clinical Trials and Hyperthermia Devices ......................................................... 710 

34.5 Conclusions ....................................................................................................................... 711 

References .................................................................................................................................... 712 


677

678 


34.1 INTRODUCTION: THE PROBLEM AND CURRENT SOLUTIONS 

A January 2006 search for the word cancer at www.clinicaltrials.gov yields a list of 4,370 federally 

and privately supported clinical trials in human volunteers. Within these results, a focus on 

chemotherapy reduces the number to 1,745 (w40% of all trials). 1 Therefore, the discovery and 

testing of chemotherapeutic agents for cancer represents a significant effort in current basic and 

clinical research. In addition, as cancer is the second leading cause of death in the United States, this 

research represents a more than significant investment in health and disease. 

Chemotherapeutics are designed to kill cancer cells or prevent them from dividing. The major 

types of cancer drugs include antimetabolites, alkylating agents, antibiotics, antimitotics, 

hormones, and inorganics (e.g., cisplatin). The National Cancer Institute (NCI) Drug Dictionary 

contains technical definitions and synonyms for more than 500 agents that are being used in the 

treatment of patients with cancer or cancer-related conditions. 1 However, the magic bullet, * interpreted 

here as “a drug that is solely specific for cancer cells,” is still not available. 2,3 The drugs that 

have been developed can kill cancer cells, but most of them can also kill normal cells, resulting in a 

low therapeutic index (ratio of lethal dose to effective dose). This limits their dosage and, ultimately, 

their efficacy. 4 Apart from the success of cisplatin in testicular cancer and methotrexate in 

choriocarcinoma, systemic chemotherapy has not yet exposed solid tumors to sufficiently high drug 

concentrations for an adequate period of time to cause meaningful tumor regressions. 4 Therefore, 

surgery † and radiation are the mainstays for local control of tumors. ‡ In modern clinical practice, 

chemotherapy is mostly used as a post-surgical adjuvant to remove microscopic metastases that 

have spread from the original site. Other applications include neoadjuvant therapy to downsize 

local tumors before surgery or radiation and palliative care for the incurable patient. 4 

One solution to the low therapeutic index of these non-specific chemotherapeutic drugs is to 

target the drugs to solid tumors by associating them with microscopic or submicroscopic carriers 

that possess some affinity for the tumors. Microscopic particles such as microspheres accumulate in 

tumor capillaries by chemoembolism, whereas submicroscopic carriers such as liposomes accumulate 

perivascularly through passive extravasation and/or active targeting to cancer cells. 5–10 It is 

here, in the context of drug delivery, that nanotechnology has already been explored for over 40 

years as an approach for reducing toxicity and increasing the accumulation of drugs in tumors when 

compared with free drug. As depicted in Figure 34.1, attention has principally focused on either 

self-assembling systems such as polymeric micelles formed from amphiphilic block co-polymers, 

11–14 phospholipid-based liposomes, 15,16 and polymer surfactant polymersomes; 17 or 

covalent-linked structures such as polymer–drug conjugates, 18–21 dendrimers, 22,23 and radioactive 

antibodies. 24,25 

The primary goal of nanotechnology for cancer therapy is to change the biodistribution of 

drugs, reducing free drug toxicity and favoring tumor accumulation rather than normal tissue 

accumulation or excretion. In order to start to achieve this goal, a nanoscopic formulation must 

satisfy the well-known requirements of a drug delivery system: load and retain the drug in or on a 

submicroscopic or nanoscopic carrier to prevent extravasation in healthy organs and tissues (e.g., 

heart and kidneys for doxorubicin) 26,27 and thereby reduce toxicity compared to free drug administration; 

28,29 evade the body’s defenses (the stealth effect) and avoid premature uptake into the 

reticuloendothelial system (RES), making the carrier available in the blood stream; 30 and achieve 

carrier uptake in solid tumors by the enhanced permeability and retention effect (EPR effect or 

* 

If an organism is pictured as infected by a certain species of bacterium, researchers can hope to discover a drug with a 

specific affinity for these bacteria and no affinity for the normal constituents of the body. Such a substance would then be a 

magic bullet. 

† 

About 90% of cancer patients undergo some kind of surgery, but only about 13% of cancers can be cured by surgery alone. 

‡ 

Although safety has been improved, people are still dying of the side effects of radiation given for cancer such as Mo 

Mowlan, former Northern Ireland Secretary, Aug 19th, 2005. 


Nanoscale Drug Delivery Vehicles for Solid Tumors 679 

Hydrophilic 

Drug 

Lipophilic drug 

Targeting 

antibody 

Targeting 

group 

drug 

passive targeting) that is based on the inherent leakiness of tumor microvessels. 31,32 This accumulation 

may possibly be aided by active targeting using antibodies. 32,33 

A recent review by Allen and Cullis outlines many of the problems that have been addressed 

and solved for nanoscale carrier systems, and it counts the successes for 13 systems that have 

gained FDA approval. If the focus is only on the anti-cancer applications, the systems, manufacturers, 

indications, and years of approval are: 8 

1. The anti-CD20 radioactive antibodies for non-Hodgkin’s lymphoma (NHL), 90 Y-ibritumomab 

tiuxetan (Zevalin by Biogen IDEC in 2002) 34 and 131 I-tositumomab (Bexxar by 

Corixa and GlaxoSmithKline in 2003). 

2. An anti-CD33-linked calicheamicin (Mylotarg by Wyeth-Ayerst) for CD33 C relapsed 

acute myeloid leukemia (2002). 

3. The three liposomal products 

a. Stealth (PEG * -stabilized) liposomal doxorubicin (Doxil/Caelyx by ALZA and 

Schering Plough) approved for Kaposi’s Sarcoma (1995), refractory ovarian cancer 

(1999), and refractory breast cancer (Europe 2003 for patients with cardiac risk) 

b. Liposomal daunorubicin (DaunoXome by NeXstar Pharmaceuticals and Gilead) for 

advanced, HIV-associated Kaposi’s Sarcoma (1996) 

c. Liposomal doxorubicin (Myocet by Elan) for metastatic breast cancer in combination 

with cyclophosphamide (Europe 2000) 

4. Adjuncts including PEG–granulocyte colony stimulating factor (pegfilgrastim/Neulasta 

by Amgen) for reduction of febrile neutropenia associated with chemotherapy (2002) 

In further depth, how are some of these nanotech-formulated drugs faring in terms of clinical 

use and outcome? For the anti-CD20 antibody Zevalin, a randomized, controlled phase III trial 

supported its accelerated FDA approval for NHL. 35 This study included 143 subjects with relapsed 

or refractory, low grade or follicular NHL, or transformed B-cell NHL. An overall response rate of 

80% was obtained in subjects receiving the Zevalin therapeutic regimen (73 subjects) compared to 

56% for subjects receiving Rituxan alone (70 subjects). Complete response rates (with no detectable 

cancer) were 30% for the Zevalin group and 16% for Rituxan alone. The side effects of Zevalin 

drug 

drug 

drug drug 

drug 

drug 

drug drug 

Encapsulated Drug 

Targeting Group 

(a) (b) (c) (d) (e) 

FIGURE 34.1 Schematic illustrations of the most developed nanoparticles for drug delivery: (a) polymer 

micelles; (b) PEGylated liposome showing encapsulated hydrophilic drug, membrane incorporated hydrophobic 

drug, and a targeting antibody; (c) a polymer drug conjugate that can also be targeted; (d) a dendrimer with an 

encapsulated drug as well as targeting ligands; and (e) a radioactive antibody. 

* PEG, Poly(ethylene glycol). 


680 


include thrombocytopenia and neutropenia. This bone marrow suppression, sometimes persistent in 

nature, has led to severe hemorrhaging and infection in a small number of patients. 35,36 Since its 

approval in 2002, further studies have supported the efficacy of Zevalin for NHL with response 

rates greater than 80%; however, the mean time to progression for these patients has still been 

limited to only 9–12 months. 25,37–39 

With over 40 years of research dedicated to liposomes alone and over 20,000 papers in the 

literature, the question can be asked: has liposomal drug delivery been successful? The first 

liposomal carrier, AmBisome (liposomal amphotericin B), was developed for severe fungal infections 

40 with a potential market of approximately 100,000 cases per year in the United States. 

AmBisome was originally approved in Europe in 1990 (Gilead and NeXstar Pharmaceuticals), 

and it was cleared for use in the United States in 1997 (Gilead and Astellas). It is now available in 

more than 44 countries worldwide. 41 This formulation then is one of the first successes in therapeutic 

nanotechnology. Clinical trials are now underway to examine the expanded use of 

AmBisome for prophylaxis against systemic fungal infection following bone marrow transplantation 

and other anti-cancer treatments. However, as discussed by Becker et al., although the 

efficacy of high-dose AmBisome is equal to or better than free amphotericin B, failure rates still 

give cause for concern. 42 One explanation for this might be the limited immediate bioavailability of 

amphotericin B when it is liposome-encapsulated. This is a good example of a common problem 

facing all nanotech drug delivery systems; that is, the challenge of reducing toxicity of free drug is 

solved, but the flip side is compromised bioavailability. 

As for liposomes against cancer, the FDA has approved two anthracycline formulations, 

Daunoxome (liposomal daunorubicin) and Doxil (PEGylated liposomal doxorubicin). These 

drugs were originally tested and approved for Kaposi’s Sarcoma, an AIDS-related malignancy 

with an incidence of only 0.02%–0.05% in the U.S. 43 Doxil has also recently been indicated for 

refractory ovarian cancer that accounts for up to 70% of the 25,500 annual cases of ovarian cancer 

in the U.S. 44 Response rates of 12%–20% have been reported for Doxil in this patient group, 

proving it to be one of the most effective salvage therapies for ovarian cancer patients, especially 

for those that have developed platinum resistance. 45,46 However, complete responses are rare, and 

chronic recurrence leads to an average survival time of only 12–24 months after therapy. 44,47 One 

of Doxil’s successes is that it is associated with lower incidences of cardiotoxicity, bone marrow 

suppression, and nausea than free doxorubicin, but it can cause other side effects, including handfoot 

syndrome and stomatitis. 28,48,49 Aside from the formulations currently approved, are new 

liposomes in development or are there new indications for previously approved liposomes? A 

search for cancer, chemotherapy, and liposome at www.clinicaltrials.gov produced 25 currently 

recruiting trials, mainly testing liposomal doxorubicin in combination with other free drugs for a 

variety of cancers as listed in Table 34.1. 50 

Liposomes were discovered in 1965 51 and explored almost immediately in applications for 

drug delivery, principally for cancer. 52 They now represent the most advanced nanoscale drug 

delivery system in clinical use. Apart from radioactive antibodies, * other nanoscale molecules and 

particles such as polymer–drug conjugates and polymer–drug micelles are being developed but 

have not yet progressed to clinical trials. However, even the liposomal nanotechnology effort 

accounts for only 25/4370 or 0.6% of all cancer-related clinical trials. This review of the literature 

and current clinical trials is illuminating and shows that there is a huge opportunity for nanotechnology 

in the form of nanomedicines to impact cancer therapy. If successful (and failed) 

designs, tests, and trials are collectively built upon, existing nanotech systems may be integrated 

with new nanotech ideas may be integrated to achieve more optimal outcomes for the delivery of 

non-specific drugs—that is, a reduced toxicity compared to free drug and an enhanced accumulation 

and bioavailability within the tumor. It is here that the thermal sensitive liposome (LTSL) 

* A search for monoclonal, antibody, and cancer found 297 trials, but only one phase I, using radio-labeled antibodies. 



TABLE 34.1 

Cancers Under Investigation for Therapy with Liposomal 

Doxorubicin 

Breast cancer 

Cervical cancer 

Endometrial cancer 

Male breast cancer 

Ovarian epithelial cancer 

Ovarian germ cell tumor 

Fallopian tube cancer 

Peritoneal cavity cancer 

Prostate cancer 

Non-small cell lung cancer 

Small cell lung cancer 

Adult hepatocellular carcinoma 

Cutaneous T-cell lymphoma 

Adult NHL 

Mycosis fungoides/sezary syndrome 

Adult Hodgkin’s lymphoma 

Lymphocyte depleted 

Mixed cellularity 

Nodular sclerosing 

Adult large cell lymphoma 

Anaplastic 

Recurrent diffuse 

Recurrent immunoblastic 

Multiple myeloma 

offers a new paradigm for local drug delivery and could provide the motivation and experience for 

other nanotechnologies that not only encapsulate drugs but also release them, making them 

bioavailable at the tumor site. 

This chapter will compare and contrast a number of contemporary liposome formulations 

(e.g., Doxil, Myocet, DaunoXome, and the currently trialed Onco-TCS). In particular, it will 

discuss the rationale and events that have led to the phase I/II testing of the doxorubicincontaining, 

temperature-sensitive liposome (ThermoDox) in prostate, breast, and liver cancer. 

This a particular example of how old and new concepts, especially ones that are bio-inspired, 

may be combined to face the challenges of cancer nanotherapeutics. The new capsular design 

addresses an unmet need in liposomal drug delivery, that of getting the drug out of the 

nanocarrier and thereby improving bioavailability. 53 This temperature sensitive liposome formulation 

containing lysolipid in the membrane (designated in the literature as LTSL) has been 

specifically designed to rapidly release drugs in combination with mild hyperthermic temperatures. 

In preclinical trials, the LTSL successfully delivers doxorubicin to solid tumors in 

bioavailable quantities that cannot be achieved by free drug administration or passive liposome 

accumulation, even with stealth liposomes. As depicted in Figure 34.2, this formulation is 

establishing a new paradigm for localized drug delivery to tumor tissue—one that moves 

nanotechnology for cancer therapy from the requirements of drug retention and carrier extravasation 

to the accomplishment of drug release in the blood stream of the tumor, thereby 

making the drug bioavailable. This gives the drug access to endothelial cells as well as 

neoplastic cells, offering the potential for anti-vascular as well as anti-cancer drug therapies. 54 


682 

FIGURE 34.2 (a) A new paradigm for drug delivery, i.e., drug is released from circulating temperaturetriggered 

liposomes in the blood stream of the tumor; drug gains access to tumor endothelia as well as to the 

neoplasm. (b and c) Video micrograph images of tumor vasculature before (b) and 24 h after (c) treatment with 

temperature-sensitive liposomes containing doxorubicin (Dox–LTSL) with an almost complete elimination of 

functioning blood vessels. (From Kong, G., et al., Cancer Res., 60 (24), 6950–6957, 2000; Chen, Q., et al., Mol. 

Cancer Ther., 3 (10), 1311–1317, 2004. With permission.) 

34.2 THE CHALLENGES FACING NANOTHERAPEUTIC INTERVENTION: 

A BRIEF REVIEW OF LIPID-BASED NANOCARRIERS 

34.2.1 GETTING THE DRUG IN 

Small molecules are on the order of a few nanometers. A liposomal carrier by definition must be 

larger than the drug it is carrying. Furthermore, a larger liposome can encapsulate a much greater 

volume, resulting in a higher drug to lipid ratio (for water-soluble drugs). * This is important 

because an overload of carrier material can impair liver metabolism and phagocytosis (RES 

uptake) and can cause granulomatous inflammation of the liver. In fact, large doses of empty 

liposomes were used in the past to block the RES and promote longer liposome-circulation halflives. 

55,56 Figure 34.3 shows that although only two molecules thick (w5 nm), the lipid bilayer † 

becomes a significant fraction of the total liposome volume at diameters of 200 nm and smaller. For 

reasons to be later discussed, liposomes are traditionally w100 nm in diameter. At this size, the 

membrane already takes up 30% of the total liposome volume. Because of mechanical restrictions 

imposed by membrane thickness and bending stiffness of the lipid bilayer itself (limited curvature), 

the smallest possible liposome has a diameter of w30 nm, exhibiting a hydrophobic volume 

fraction of w80% and a limited capacity to carry water soluble drugs 

Therefore, for these nanoscale carriers, how much drug could be loaded in a 100 nm capsule? 

The number of molecules of hydrophilic drug per liposome (N d) depends on the solubility (S d)of 

the drug as well as the internal radius (Ri) of the liposome 

N d Z 4 

3 pR3 i S dA 0 

Note that Avogadro’s number is included in the calculation. For cisplatin with a solubility of 

5.8 mM at 258C, a 100 nm liposome is capable of encapsulating only 1,340 molecules of this 

drug. By comparison, the 100 nm liposome is composed of about w140,000 molecules of lipid, 

giving a drug to lipid ratio of only 0.005 (w/w). 

The drug doxorubicin, on the other hand, is amphiphilic with a protonatable amine (pKA 8.4); 57 

therefore, depending on the pH, a fraction of the drug is soluble in the bilayer. This fraction can 

transit the bilayer if an electrochemical gradient is imposed, facilitating the active loading method 

* 

This chapter does not consider hydrophobic drugs. 

† 

The hydrophobic bilayer thickness is w4 nm. 




Volume fraction 

1 

0.9 

0.8 

0.7 

0.6 

0.5 

0.4 

0.3 

0.2 

0.1 

0 

10 

30 

50 

70 

90 

110 

130 

that has revolutionized liposomal drug loading technology for weak acids (Figure 34.4). 58,59 This 

method results in extensive loading of up to 61,000 molecules of doxorubicin per liposome (drug to 

lipid ratio of 0.3) with the formation of a complex of drug crystals inside the liposome. In considering 

new nanoscale drug delivery systems, these kinds of simple calculations are necessary to 

determine carrying capacity and drug dosage. As previously mentioned and will be later discussed, 

nanoscopic carriers are readily identified by the immune system as foreign and can become toxic to 

the RES in high doses. There is, therefore, a delicate balance between efficacy and toxicity, not just 

for the drug but also for the nanoscopic carrier. 

150 

170 

Outer liposome diameter (nm) 

190 

Hydrophobic 

Aqueous/ 

Hydrophilic 

FIGURE 34.3 Plot of hydrophobic volume fraction as a function of liposome diameter for single bilayer 

vesicles. For the traditional 100 nm diameter liposome, 70% of the liposome volume is available for aqueous 

content. However, for the smallest liposome possible (30 nm because of limited curvature), there is considerably 

more hydrophobic volume per liposome (w80%). 

Dox-H + /Citrate fiber 

Bundle complex 

2 

Citrate 

Inner surface 

outer surface 

Citrate 

H + 

Dox-H + 

Dox o 

H + 

Dox-H + 

Dox o 

Dox-H + 

1 

Dox o 

H + 

Dox-H + 

FIGURE 34.4 Remote loading of doxorubicin into liposomes using a pH gradient. Doxorubicin, an amphiphilic 

drug, crosses the lipid bilayer in neutral form. When it encounters the acidic interior of the liposome, an 

amine is protonated to form Dox-HC. This compound is retained on the interior, forming fiber bundles by 

electrostatic interaction with citrate ions. 


Dox o 

H +

684 

34.2.2 GETTING THE NANOCAPSULES TO THE TUMOR 


The next issue to address for any nanocarrier is the probability that it will reach the tumor microvessels 

after injection into the blood stream. Considerations here include circulation through the 

tumor and pharmacokinetics of the carrier. A large fraction of the human cardiac output circulates 

through the kidneys and liver, an this explains the rapid clearance of particles by these organs. It is 

estimated that only a small fraction of the cardiac output actually goes to the peripheral circulation 

(and to many tumors). 60 Within this blood, what concentration of liposomal drug is actually 

present? Liposome extrusion methods are limited to a lipid concentration of 100 mM, above 

which the high viscosity of the lipid suspension limits processing. Assuming a clinical injection 

volume of 20 mLs, the lipids would be diluted 250 times (in 5 L blood) to a plasma concentration of 

0.4 mM lipid. If drugs such as doxorubicin can be loaded into liposomes at internal drug concentrations 

of 300 mM, then the actual plasma drug concentration (still encapsulated) is now only 

0.216 mM. This is about 200 times the amount required for a log kill of cells in culture (w1 mM 

Dox). However, this assumes complete mixing and no clearance (see below). As the complexities 

of drug transport become apparent, it is clear that delivery (and not just reduction of toxicity) is the 

major goal of nanotherapeutics. 

Given these limits on actual blood levels available for delivery, how does the nanoparticle 

accumulate in the tumor? Where does it actually collect within the tumor, and does it stay there? 

Either the carrier must lodge in the microvascular bed of the tumor as in chemoembolism strategies, 

or it must leak out of the tumor vasculature (by the EPR effect) as in most nanoparticle strategies. 

The vascular permeability of a drug (across the endothelial lining) is limited by the size and charge 

of the drug as well as the structures of the vessel wall, particularly the basement membrane. 61 For 

small therapeutic drugs such as doxorubicin, permeabilities through normal vasculature are on the 

order of 10 K4 cm/s. Similarly, sulforhodamine has been measured to have a permeability of 3.4! 

10 K5 cm/s. 62 As the size of a nanoparticle increases beyond 6 nm, its ability to extravasate through 

the endothelial lining drops precipitously. 63 For example, the microvascular permeability of 

albumin (radius of gyration 7 nm) in normal tissues is 25!10 K8 cm/s, and that for 150,000 MW 

dextran (radius of gyration 11 nm) is 7.3!10 K8 cm/s. 62,64–66 Therefore, these polymers and 

proteins are well retained in the blood. For 100 nm liposomes, the normal vascular permeability 

is also limited at 9!10 K8 cm/s. 65 Therefore, liposomes are over 1,000 times less permeable across 

normal vessels than free doxorubicin. It is this physical barrier that is actually responsible for the 

reduction in toxicity, dramatically altered biodistribution, and somewhat extended circulation halflife 

of liposome-encapsulated drug compared to free drug. For example, conventional liposomal 

daunorubicin (Daunoxome) is retained in the blood with a circulation half-life of 4 h compared to 

0.8 h for free daunorubicin. 67 As the size of the nanoparticle is increased, motivated by the need to 

create a large encapsulated volume, the liposome gains the unique ability to keep the drug in the 

bloodstream and away from healthy organs and tissues such as the heart and kidneys. This was one 

of the major successes of liposomes, a clear demonstration that nanotechnology could impact the 

therapeutic toxicity of non-specific drugs like doxorubicin. 28,48,49 In fact, improved toxicity profiles 

have been the primary reason for approval of conventional liposome formulations like Daunoxome 

(FDA 1996) and Myocet (EU Commission 2000). 

But how does this aid in the delivery of drugs to tumors? Fortunately, as already intimated, 

tumors have leaky blood vessels that are more permeable than normal blood vessels. For example, the 

tumor microvascular permeability for 150 kDa dextran is 5.7!10 K7 cm/s, eight times higher than in 

normal tissue, 64 and for albumin, tumor vessel permeability is 25 times higher (1!10 K7 cm/s). 66 It 

has been shown that this permeability extends to nanoparticles such as colloidal carbon 68 and liposomes 

up to 400 nm in diameter. 69 Therefore, one favored mechanism of targeting tumors has been to 

take advantage of the natural leakiness of tumor blood vessels to certain nanoparticles such as 

drug–polymer complexes and liposomes. 9,70,71 These drug delivery systems achieve localization 

simply by passive accumulation. Dosimetry results for the conventional liposome Daunoxome show 



that approximately 2.84 ng/mg (5 mM) of daunorubicin accumulates in the tumor in the first 24 h 

after an injected dose of 2 mg/kg body weight. 68 This concentration of drug is actually fairly 

successful at killing tumor cells in culture (IC50 2.8 mg/ml). 72 So why does this formulation 

show limited success in the clinic? The problem is probably not just inadequate liposome 

accumulation but also the inability to make the drug bioavailable in a fast enough time period, as 

will be discussed later. 

Returning to a discussion of dosimetry, an important challenge was to determine what limits 

liposome accumulation in the tumor and how this uptake may be increased. It was soon discovered 

that the reason for limited tumor uptake is simply the limited circulation half-life of conventional 

liposomes (2–4 h). 73 These liposomes are labeled by complement and other opsonins as foreign and 

are subsequently removed by the phagocytic RES of the liver and spleen. 74,75 Incidentally, particles 

larger than 0.6 mm are cleared by the spleen regardless of opsonization. 76 This is another major 

lesson for any nanotechnology-based drug carrier system—that the surface chemistry of particles 

(especially nanoparticles) is crucial in the foreign-body reaction. Essentially, the early researchers 

and companies (The Liposome Company/Elan and Nexstar), intending to target cancer, unwittingly 

created an artificial parasite. Some formulations, however, have cleverly taken advantage of this 

natural localization to the Kupffer cells of the liver for treatment of the parasite Leishmania that 

resides in these same cells. 77–79 By creating a highly immunogenic (opsonizable) liposome 

containing negatively charged lipids, these formulations were rapidly removed from circulation 

by the RES. A similar liver-targeting technique was put into clinical practice when Ambisome was 

approved for use against leishmaniasis. 80,81 

The solution to the clearance problem was to favorably modify the liposome surface chemistry. 

Inspired by the very long (120 day) half-life of the red blood cell, Allen and Papahadjopoulos 75 

explored the idea of incorporating GM1, a red blood cell membrane glycolipid, into the liposomal 

membranes (10 mol%). This conferred longevity to the liposomes, increasing plasma half life to 

w24 h. However, GM1 is an expensive molecule to purify from whole blood. Therefore, they 82 and 

others 83–86 turned to the chemical synthesis and development of polyethylene glycol derivatized lipid 

(e.g., DSPE–PEG * ) that similarly increased circulation time. This conjugated lipid had a single 

negative charge because the derivatization left the normal zwitterionic PE charged at the phosphate. 

Although the incorporation of the negatively charged PS was highly immunogenic, their original 

hypothesis was that when incorporated into the bilayer at a few mol%, this shielded negative charge 

(similar to that of GM1) was the source of resistance to opsonization (Lasic, D., Personal Communication. 

Liposome Technology, Inc., Menlo Park, CA, 1992). Subsequently, evidence has been 

accumulated that indicates that the mechanism is primarily one of steric stability because the electrostatic 

potential lies somewhat underneath the steric barrier profile provided by PEG as demonstrated by 

McIntosh et al. 87,88 The mechanism for the long circulation lifetime conferred by incorporating GM1 is 

still unknown, although Huang postulated a dysopsonization for this ninja liposome. 89,90 

Preclinical animal studies demonstrated the dramatic changes in circulation half-life, tumor 

accumulation, and anti-tumor efficacy that resulted from these somewhat bioinspired surface 

chemical strategies. 30,82,91–96 The longer circulation time of PEG-grafted lipids allowed for 

increased extravasation from leaky tumor microvessels, successfully exhibiting the EPR paradigm 

of passive targeting. The new so-called “Stealth” technology was developed by Sequus and incorporated 

into a PEGylated liposomal doxorubicin formulation, Doxil, by Johnson and Johnson. 

Doxil has achieved circulation half lives of 20 h in mice 82 and w50 h in humans. 97 Furthermore, 

at 48 h after injection, the median tumor concentration of Doxil (in Kaposi’s sarcoma lesions in 

humans) was 19 times higher than in normal skin. 97 In mice, about 7 ng/mg of Dox accumulated in 

FaDu xenografts (human squamous cell carcinoma) after injection of 5 mg/kg Doxil 

(Figure 34.24). 98 

* DSPE–PEG, Distearoylphosphatidylcholine grafted with polyethylene glycol. 


686 

Expansion modulus K A mN/m 

4000 

3500 

3000 

2500 

2000 

1500 

1000 

500 

0 

0 20 40 60 80 100 120 140 

t 1/2 (min) 

FIGURE 34.5 Plasma half life of conventional liposomes versus elastic modulus of the membrane (for normal 

RES). Membranes with a higher elastic modulus circulate for longer times. That is, a low resistance to 

mechanical expansion of the bilayer interface is correlated with liposome destruction in the blood stream. 

This points to a mechanical mechanism for the presumed opsonization of liposomes. A similar trend holds 

when the RES is pre-saturated with liposomes, but half-lives are on the order of hours. 

An alternative mechanism to resist opsonization that also relies on reducing protein adsorption 

and binding at the bilayer head group region was to modify the membrane mechanically rather than 

to chemically modify the surface. This was achieved by using a lipid composition that creates a very 

tight interface among the headgroups as measured by a high area dilation modulus for the bilayer. 99 

This composition utilizes lipids with saturated acyl chains and incorporates high concentrations 

(w50 mol%) of cholesterol in the membranes. Taking data from the literature on circulation half 

lives and comparing them to the elastic area dilation modulus measured using the micropipette 

technique for similar lipid compositions, a direct correlation was found between circulation half life 

and bilayer elasticity as shown in Figure 34.5. 87 

This shows that resistance to bilayer area expansion is a key property in achieving a long 

circulation performance for these nanocarriers. That is, a high resistance to mechanical expansion 

of the bilayer interface is correlated with reduced opsonization in the blood stream. This points to a 

nanomechanical mechanism for the presumed opsonization of liposomes where the molecular 

insertion that underlies opsonization must create area expansion in the bilayer. A similar trend 

holds when the RES is pre-saturated with liposomes. Here, half-lives are on the order of hours, and 

clearance may be due to destruction in the blood stream. Again, low membrane compliance resists 

this and is the property that correlates with long circulation performance. 

A particular example is the sphingomyelin/cholesterol bilayer with a high area dilation 

modulus of w2,200 mN/m 99 that has been utilized in Onco-TCS, the liposomal vincristine * 

formulation of Inex. 100 Onco-TCS achieves a prolonged circulation half life (6.6 h) when 

compared to free vincristine (1.36 h) in preclinical studies. 101 Therefore, a low bilayer membrane 

compliance can be used to increase circulation time. Onco-TCS also exhibits a higher concentration 

in tumors and demonstrates less neurotoxicity and greater efficacy in mice bearing L-1210 

and P-388 leukemias. This formulation has been shown to have a similar MTD when compared 

with free vincristine, and it is currently in phase II/III clinical trials for aggressive recurrent 

NHL. 8,102–104 

* First extracted from the Madagascar periwinkle. 




34.2.3 MECHANISMS OF LIPOSOMAL DRUG DELIVERY 

As the study of liposomes continued, it was clear that in order to address mechanistic questions 

about drug delivery such as the EPR effect and the distribution of liposomes and other nanoparticles 

within the tumor, a direct in vivo measurement in real time was required. The skin flap tumor 

window chamber, first used in the 1940s for observation of tumor microvasculature, offered this 

capacity. 105 This research team 54,71,106 and others 9,66 have extensively utilized the window 

chamber to measure liposomal circulation half-life, accumulation in tumors, and tumor distribution 

both with and without hyperthermia. For the conventional, stealth, and temperature-sensitive liposomes, 

this technique has been invaluable in the direct observation of accumulation and/or 

triggered release of encapsulated drug in the tumor tissue. The window chamber is shown in 

Figure 34.6. 

The procedure involves implanting tumor cells onto the exposed fascia and covering the 

surgical site with a glass window. After the tumor grows, the vasculature is visible by intravital 

microscopy with transmitted white light. In addition, using epifluorescence, it is possible to quantitatively 

evaluate fluorescent or fluorescently labeled molecules within the vasculature and within 

the tumor tissue over time. With this model, complex drug delivery data such as pharmacokinetics, 

vessel wall permeability, and intratumoral distribution have been quantified. 105 Using small fluorescent 

molecules fluorescently labeled albumin, 62 and fluorescently labeled liposomes, 65 the 

following have been shown: 

1. Small molecules such as sulforhodamine (606 Da) and proteins such as albumin 

(66,430 Da) are readily extravasated into tumor tissue, supporting the established permeability 

of tumor vasculature. 62 

2. Conventional liposomes also extravasate but the extent of accumulation in the perivascular 

space is severely limited by fast RES uptake and disappearance from circulation 

(Figure 34.7, middle panels and right graph). 65 

3. Stealth liposomes have an extended circulation half-life and accumulate to a greater 

extent as shown in the epifluorescent images of the tumor vasculature in Figure 34.7 

(left panel and left graph). 65 This supports the early efficacy studies by Vaage and 

FIGURE 34.6 (See color insert following page 522.) Schematic of rat with a window chamber surgically 

attached to the dorsal skin flap (left). The skin flap is a 200 mm thick section of dermis with implanted tumor 

cells. As these cells grow, the tumor is visible via trans-illumination (right) or epi-illumination. This model was 

used to quantify complex data such as tumor vascular permeability and liposome accumulation. (From Huang, 

Q., et al., Nat. Biotechnol., 17, 1033–1035, 1999. With permission.) 


688 

(x) 90 min after injection' 1 min after injection Transillumination 

(a) (b) 

100 DOXIL-Like 

100 

Normalized liposome amount 

80 

60 

40 

20 

Stealth in tumor 

0 0 15 30 45 60 75 90 

Time (min) 

Conventional in tumor 

(a) (b) (c) 

(d) (e) (f) 

(g) (h) (i) 

Vascular 

Interstitial 

Normalized liposome amount 

Metcalfe that showed that stealth, but not conventional, liposomes produced tumor 

efficacy in mouse tumor models. 93–96 

What this data 99,107,108 clearly shows is that the stealth liposomes out-perform the 

conventional liposomes with respect to circulation half-life and accumulation in 

the heterogeneously leaky tumor (left and center panels). Furthermore, as shown in the 

right panel, normal blood vessels (in the same window chamber) maintain a seal against 

liposome extravasation. The small 100 nm diameter of the liposome and the incorporation 

of PEG-lipid (4–5 mol%) work to preferentially increase the amount of encapsulated 

80 

60 

40 

20 


Stealth in normal 

EVACET-Like 

0 0 15 30 45 60 75 90 

Time (min) 

Vascular 

Interstitial 

FIGURE 34.7 Video micrographs showing the window chamber vasculature via trans-illumination (top row) 

and epi-illumination (second and third rows) after injection of stealth liposomes in tumor (left column), 

conventional liposomes in tumor (middle), or stealth liposomes in normal tissue (right). The second row 

shows 1 min after injection, and the third row shows 90 min after injection. The graphs show the relative 

amounts of liposomes in the blood stream and in the tumor versus time (stealth on left and conventional on 

right) as calculated from the fluorescent images. (From Wu, N. Z., et al., Cancer Res., 53 (16), 3765–3770, 




FIGURE 34.8 Schematic showing the traditional paradigm for liposomal (and polymer) drug delivery. In this 

paradigm, extended circulation time takes advantage of leaky tumor vasculature to preferentially deliver 

carriers to the tumor tissue. 

drug being delivered to the tumor while normal tissue is protected. This behavior 

establishes the current paradigm of the EPR effect for liposomes (and other macromolecules 

and nanoparticles) as depicted in Figure 34.8; that is, the basis for 

preferential drug delivery is this disparity in vascular permeability to even relatively 

large 100 nm particles, resulting in a fraction of the injected dose gaining access to the 

perivascular space of the tumor tissue. 

4. The vascular pore cut-off size for liposome extravasation in the human ovarian carcinoma 

SKOV3 was found to be between 7 and 100 nm diameter at 378C; 71 other studies have 

shown a cut-off size of up to 400 nm in the human colon adenocarcinoma LS174T. 70 

5. Mild local hyperthermia (HT) of 428C could dramatically increase the permeability of 

tumor vessels to stealth (non-thermal sensitive) liposomes up to 400 nm in diameter with 

the most dramatic effect seen at 100 nm. 71 In one experiment, the amount of liposome 

extravasation was increased by 50-fold with HT compared to normothermic controls for 

the relatively impermeable SKOV3 tumor. 109 

6. In spontaneous tumors in pet cats, HT dramatically increased the accumulation of nonthermal 

sensitive technetium-99-labeled liposomes in the tumor tissue by 2–16-fold, 

further supporting the utility of HT in spontaneous tumors. 106 

These results demonstrate the important benefits of mild hyperthermia for liposome distribution 

in the EPR paradigm (even with non-thermal sensitive liposomes). However, they also point to a 

number of flaws in relying on the EPR effect alone to deliver nanoparticles and their drug. It is 

apparent that not all tumors are permeable to nanoparticles on the scale of 100 nm. In addition, 

throughout the window chamber studies, researchers have observed that liposomes and other 

nanoparticles only accumulate in the perivascular space after extravasation. 9,71,109 Even 

antibody-targeted liposomes show only limited interstitial diffusion. 

As for the visualization of liposome distribution, Figure 34.9 shows magnified videomicrograph 

images of the microvasculature of a mammary adenocarcinoma model in a skin flap window 

chamber. 65,110 On the left, a trans-illuminated image shows blood vessels of approximately 30 mm 

in diameter. The smallest (arrow) is about 10 mm, allowing only slow passage of red cells in single 

file. On the right is an epi-illuminated image of fluorescent liposome distribution 90 min after 

injection of NBD-labeled liposomes. This characteristic hot spot pattern illustrates the gross heterogeneity 

of tumor vascular permeability and the limited transport of fluorescently labeled liposomes 

into the packed cellular bed of the tumor tissue. This image very dramatically demonstrates that 


690 


FIGURE 34.9 (a) Videomicrograph bright-field image of the microvasculature of a mammary adenocarcinoma 

in a skin flap window chamber. (b) Videomicrograph epifluorescent image of the same window chamber, 

showing a hot spot pattern of stealth liposomes 90 min after tail vein injection. The focal deposits of liposomes 

are limited to the perivascular space. (From Chen, Q., et al., Mol. Cancer Ther., 3 (10), 1311–1317, 2004. 


liposomes only accumulate in the perivascular space of the tumor and do not diffuse to cells deep 

within the tumor. In fact, the hot spots are close enough to the circulation that a portion of the 

liposomes can actually wash back into the vessels. 111 The important conclusion here is that the 

liposome-encapsulated drug does not reach the majority of the tumor cells based on the EPR effect 

alone. This is also expected for other nanoparticles that aim to deliver drugs to tumors; they will not 

reach beyond the perivascular space unless some additional chemical or external influence 

is exerted. 

Whereas distribution of nanoscopic particles only to the perivascular space is clearly a limitation, 

a bigger question is if the drug is actually bioavailable when delivered to even these regions. 

How does the drug get to its target once the carrier has accumulated in the tumor? How do the 

properties of the carrier and the local tumor environment contribute to this process? This chapter 

previously described how the currently approved liposomal nanotechnologies achieve a significant 

decrease in drug toxicity and a modest accumulation of drug in the tumor. However, the design 

requirement of retaining drug in the carrier during circulation results in most of the drug remaining 

inside the carrier if and when it localizes in the tumor tissue. It is expected that the drug will 

passively leak across the membrane, but it seems that it is often not bioavailable in high concentrations 

on the time scale necessary for therapeutic action. In particular, non-cell cycle specific 

drugs like doxorubicin should ideally be available to the intracellular targets (including DNA and 

RNA) as a sharp peak of concentration, maximizing drug effect in a short time period to prevent cell 

recovery and resistance. This inherent slow release may actually be an advantage for drugs that 

have cell cycle specificity such as the vincristine formulation Onco-TCS. 104 In this scenario, the 

slow release and long circulation half life of the sphingomyelin/cholesterol liposomes make free 

vincristine available to as many cancer cells as possible as they progress through the cell cycle to 

the G2/M stage. 

Overall, the controlled release of the drug from carriers at the tumor site is a problem that all 

polymer, liposomal, micro-particle, and capsular nanotechnologies still face. 8 For polymer–drug 

conjugates, it has been fairly straightforward to convert the covalent linkers into biodegradable 

linkers (cleaved by endogenous enzyme action or by simple bond hydrolysis). Preclinical studies 

using this method have demonstrated drug release specifically at the tumor with a resulting increase 

in anti-tumor efficacy. 112,113 However, for liposomes, as mentioned above, the main focus of 

modern carrier design has been to lessen leakage from the liposomes in the circulation phase 

rather than to trigger release at the tumor site. 



34.3 NEW CAPSULAR DESIGN FOR DELIVERY OF DRUGS TO AND RELEASE 

OF DRUGS IN LOCALIZED SOLID TUMORS 

This section will briefly review the events leading up to the invention of the temperature sensitive 

liposome (LTSL), a potential solution to this paradox of drug retention in circulation and rapid 

release at the tumor site. It will also discuss studies on the composition, structure, and property 

relationships of membranes. It was this wealth of fundamental lipid bilayer knowledge that allowed 

a liposome to be forward engineered to perform a number of functions: survive the physical and 

chemical conditions of the blood stream, accumulate in the tumor, and respond to stimuli (in this 

case, mild hyperthermia) to trigger the release of encapsulated drug. The key to the design and its 

functional performance, as in all nanotechnologies for drug delivery, is to understand the materials 

relationships of the carrier. 

34.3.1 GETTING THE DRUG OUT 

After 35 years of research, the long-circulating liposomes (e.g., Doxil) have fulfilled the requirements 

of loading and retention of drug, evasion of the body’s defenses, and accumulation in tumors 

by EPR. However, it is widely recognized that liposome performance is often hampered by the very 

lipid compositions that create low bilayer permeability and allow retention of the encapsulated drug 

during the blood-borne transport phase. The remaining goal of releasing a drug such as doxorubicin 

fast enough to kill tumor cells was not part of the original conventional (Myocet, Daunoxome) or 

stealth (Doxil) designs. One solution that now comprises the main topic for the rest of this chapter 

has been to take a materials engineering approach to liposome design and to explore nanoscale 

material defects in these 100 nm carriers. The task undertaken for this particular system was to 

characterize the relationships between lipid composition (lipids and lysolipids), bilayer nanostructure 

(grain boundary defects), and membrane properties (mechanical, thermal, and permeability) in 

order to design a liposome that releases drug rapidly upon an externally applied trigger, in this case, 

mild heat of around 428C. 114,115 

The functions (load, retain, target, and release) have been defined, and there is an established 

wealth of materials data, prototype tests, production methods, and performance studies for liposomes 

in the literature. Using these guidelines, the goal is to trigger release of the liposomes after 

extravasation by EPR. Based on size alone, liposomes greater than 600 nm are taken up by the 

spleen, and those less than 70 nm are taken up by hepatocytes. 76,116 In addition, 400 nm is the 

cut-off for tumor extravasation. 70 Therefore, the optimal liposome size for EPR-based drug delivery 

is about 100 nm. In addition, 100 nm liposomes have demonstrated the greatest increase in tumor 

uptake when combined with hyperthermia. 109 As for surface design, conventional formulations 

only provide a relatively short circulation half life (2–4 h) 73 , but PEGylated lipids with their surface 

repulsive potentials 87 tend to avoid RES uptake. Therefore, the starting point was the 100 nm 

stealth formulation, shown schematically in Figure 34.10. 

Given this liposome, researchers must remember that many physical limitations still hamper 

effective drug delivery. As discussed above, only a small fraction of the circulation passes through 

the tumor. Some tumors have low permeability, preventing uptake of even 100 nm nanoparticles. 71 

Upon extravasation, liposomes accumulate only in the perivascular space, and the liposomes that do 

accumulate achieve only slow release of bioavailable drug. In 1995, even though liposomal formulations 

were being tested and approved, the current technology was not achieving the kind of 

bioavailable drug levels necessary for increased efficacy. Through collaboration with the 

hyperthermia program at Duke Medical Center, this team found that local hyperthermia may be 

used to address many of these problems. Before discussing the materials aspects of the liposome 

that promote triggered release, the effects of hyperthermia on liposomal drug delivery must 

be considered. 


692 

Drug 

FIGURE 34.10 A tentative liposome design was a 100 nm diameter stealth liposome composed of a DPPC 

bilayer with 4–5 mol% PEG–DSPE (only shown on the outside) and encapsulating the drug doxorubicin via 

remote loading. 

The mechanisms of synergy between hyperthermia (HT) and liposomal chemotherapy are 

manifold (Figure 34.11). HT improves vascular perfusion, thereby increasing blood flow through 

tumors that usually have significant areas of poor perfusion. Furthermore, HT dramatically 

enhances transvascular permeability, increasing pore size between endothelial cells to enhance 

liposome extravasation. Tumor microvessels are particularly sensitive to the effects of HT, 

providing added selectivity that protects the normal tissues. 117,118 These changes last for several 

hours following HT treatment and have been exploited to augment liposomal drug delivery to 

tumors up to 50-fold in otherwise impermeable tumor vessels as described above. 109 In addition, 

hyperthermia provides direct cytotoxicity to tumor cells and synergism with many drugs, including 

doxorubicin and cisplatin. 117,119 A number of liposomal drugs have been tested in a variety of 

Direct 

cytotoxic 

effects 

Hyperthermia 

Synergistic 

effects with 

DOX 

Increased 

therapeutic effect 

Thermosensitive 

liposomes 

Increased 

extravasation 

Increased 

local drugh 

concentration 


Triggered 

drug release 

FIGURE 34.11 Schematic showing the mechanisms of synergism between hyperthermia and liposomal drug 

delivery. HT alone has some direct cytotoxic effects as well as synergism with many chemotherapeutic agents. 

In addition, HT enhances liposome extravasation and local tumor perfusion. The new concept was rapid 

triggered drug release from liposomes at mild hyperthermic temperatures. 



tumor models, and these reports consistently prove that HT increases liposome accumulation in 

tumors and enhances anti-tumor effect. 117 Finally, with new advances, the current HT technology 

has the ability to administer controlled local heat to superficial or deep locations with accurate 

thermal dosimetry. 118 The remaining question, then, is if the performance of stealth liposomes can 

be improved by releasing the drug using externally applied heat. The answer lies in tailoring the 

materials composition, nanostructure, and properties of the lipid bilayer itself. 

It is important to note here that the traditional target for chemotherapeutic drug delivery has 

always been distant metastases; very few primary tumors have been treated, much less cured, with 

chemotherapy alone. Therefore, these studies deviated from the traditional target for liposomes, took 

a bold and possibly unconventional track, and focused on local tumors. Whereas metastases are 

clearly the most important problem, an inability to control the local growth of solid tumors leads to 

significant morbidity and mortality in cancer patients. In the U.S., over 60% of therapies eventually 

fail with local recurrence. 120 Furthermore, improved locoregional control has been shown to enhance 

survival in many common cancers, including breast, prostate, lung, and cervical cancer. 121,122 What 

is needed here is a parenteral drug delivery system that can dramatically reduce the tumor size, to 

debulk for subsequent surgery or radiation, or maybe even achieve complete regression. 

What this team set out to do was to design, create, and test a liposome that would respond to 

clinically attainable mild hyperthermia temperatures in the range of 418C–428C. The design 

required an intravenously injected liposome that released its drug only upon entering the heated 

tumor mass. The great surprise, in the very first series of preclinical tumor growth delay studies, 

was a response of 11 out of 11 complete regressions. 123 In addition, measurements of drug accumulation 

in the tumor tissue showed concentrations three times higher than with the combination of 

Doxil-like liposomes and HT 98 in spite of similar plasma half lives. * 

34.3.2 A NEW LIPOSOME DESIGN 

The lysolipid-containing temperature-sensitive liposome (LTSL) was invented in 1996, offering an 

effective way to achieve targeted tumor drug release for cancer chemotherapy. 98,114 This formulation 

is now in phase 1 human clinical trials. 53 Preclinical studies in mice, rats, and dogs have 

shown promising efficacy, reducing tumor growth in many of the tumors tested. 50,98,123 However, 

some variations in tumor response and subtleties in mechanism are still to be investigated. As will 

be discussed at the end of this section, attempts to correlate tumor growth delay with measurable 

parameters such as hypoxic fraction, intratumoral drug concentration, and tumor drug sensitivity 

reveal a complex relationship between these parameters. The current discourse will focus on how 

this liposome achieves the performance requirement of rapid triggered release of doxorubicin at 

clinically attainable, mild hyperthermic temperatures. 

As demonstrated w35 years ago by de Kruijff 124 and Papahadjopoulos 125 and subsequently 

modeled by Mouritsen, 126 the permeability of a lipid bilayer membrane to small ions, water, and 

even small drug molecules is dramatically increased at its gel-to-liquid melting phase transition 

when compared to its gel state. At the melting phase transition temperature (Tm), gel and liquid 

domains coexist, and solid phase lipids form grain structures to be described below. At the grain 

boundaries, leaky interfacial regions are present. 126–129 The increased permeability may be attributed 

to these interfacial regions as well as the much higher lateral compressibility of the membrane 

in the phase transition region. 130,131 For HT-triggered drug release in the body, a lipid was needed 

that remained solid and impermeable at body temperature (378C) but became leaky to its encapsulated 

drug when heated just above this temperature. The lipid of choice was relatively obvious, 

DPPC † has a phase transition temperature of 41.98C 132 and so has been used in many studies that 

aimed to explore temperature-sensitive drug delivery. 117,133 

* 

The LTSL, if anything, showed a shorter half-life. 

† 

DPPC, Dipalmitoylphosphatidylcholine. 


694 

Thermal sensitive liposomes (TSLs) were not new in 1996. Previous formulations composed of 

DPPC mixed with other lipids like DSPC * , 133 DPPG † , 134 or HSPC ‡ :Cholesterol, 135 all required a 

temperature of 438C–458C to reach the peak in membrane permeability. The two- (or more) 

component phase diagrams show the origin of this transition temperature elevation, an ideal 

thermodynamics of mixing (entropic) that broadens the transition and raises it compared to the 

pure DPPC bilayer. 136–138 As a result, these formulations were not well matched to the clinical 

setting as hyperthermia temperatures greater than 438C are not clinically attainable and may cause 

discomfort and vascular damage to the patient. 118 Also, even at the transition temperature, the rates 

of drug release for these lipid mixtures were still relatively slow (15–60 min). This is much longer 

than the transit time of a liposome through the tumor circulation, so these early TSLs relied on the 

EPR effect to position the liposomes in the tumor before releasing the drug. Unfortunately, the 

stealth technology had not yet been invented when these TSLs were being tested, and they did not 

accumulate in the tumor enough to be efficacious. Paradoxically, the stealth liposomes achieved 

accumulation but did not release their drug. Because many development agendas for conventional 

and stealth liposomes were already in place, TSLs received less attention for many years. With a 

renewed interest in thermally triggered systems and with the benefit of hindsight, this team began 

the LTSL design in 1996 by combining old and new technologies (thermal sensitivity and stealth 

PEG). In addition, experience from many years of studying, measuring, and modeling the material 

properties of lipid bilayers using micropipette manipulation techniques was applied. 

The micropipette manipulation system has been developed and used by researchers to measure 

the mechanical properties of lipid bilayer vesicles, 99,107,108 permeability of lipid membranes, 139–141 

molecular exchange with lysolipids, 142 and other mechanical and thermomechanical properties 

143,144 in order to characterize bimolecular membranes as previously reviewed. 145–147 The 

micromechanical experiments on lipid vesicles established several key findings that influenced 

the new design. 

1. Experiments that measured the resistance of lipid vesicle bilayers to area expansion and 

so derived the bilayer elastic expansion modulus showed that liquid state (La) lipids such 

as egg PCs plus cholesterol provided sufficient strength to retain drug and remain intact 

in the blood stream. 

2. The more saturated PCs § and sphingomyelin plus cholesterol exhibited enormous elastic 

moduli approaching traditional plastics 99 and were even more impermeable to water 139 

and especially drugs. 

3. The inclusion of cholesterol (to achieve strength and impermeability) abolished the 

melting phase transition and so attenuated any possible thermal sensitivity for DPPC 

(e.g., DPPC:Cholesterol in a molar ratio of 1:1). 135 

4. Even non-cholesterol-containing PC bilayers had moderate strength as gel phase 

membranes that were essentially thin (5 nm) layers of wax. 107 

Therefore, a liposome using gel phase lipids was designed. The design was built around DPPC 

with its T m of 41.98C plus a PEG -conjugated lipid (DSPE–PEG). The inclusion of DSPE–PEG did 

not change the transition permeability compared to pure DPPC. 148 Ideally, a liposome with a 

slightly lower Tm and higher membrane permeability than that of pure DPPC is desired. The 

additional component that achieved these goals will be discussed later in this section. First, the 

* 

DSPC, Distearoylphosphatidylcholine. 

† 

DPPG, Dipalmitoylphosphatidylglycerol. 

‡ 

HSPC, Hydrogenated Soy Phosphatidylcholine. 

§ Phosphatidylcholines. 

Poly(ethylene glycol) 




structural features that apparently made the liposomes leak their drug, i.e., the grains and grain 

boundaries, must be discussed. 

34.3.3 NANOSTRUCTURES AND MEMBRANE PERMEABILITY 

In 1987, working with giant lipid vesicles (20–30 mm diameter), this team showed that a single lipid 

vesicle could be heated or cooled through its main phase transition at the tip of a micropipette. 107,143 

This allowed the transition to be characterized in terms of the bilayer area change reflecting a 

w25% change in cumulative area per molecule (from 65A 2 in the liquid phase to w41A 2 in the Pb 0 

solid phase). Under no applied membrane tension, the solid vesicle was crumpled and faceted, 

whereas under controlled low tension (suction), it formed flat plate-like domains. Upon melting, 

these lipid 2D icebergs flowed past the pipet tip as they moved in a sea of melting lipid. In the gel 

(solid) state, it was shown that the rippled nanosuperstructure that characterized the P b 0 phase could 

be pulled flat. 143 Then, subjecting the vesicle to shear deformation, a yield shear and shear viscosity 

for the membrane were measured. The values for these material properties were on the order of 

simple waxes and plastics just below their melting temperatures. The membrane behaved like a 

Bingham plastic, yielding after an elastic shear deformation and then flowing with a velocity that 

was proportional to the excess stress (provided by the micropipette suction pressure) above the 

yield point. Both the yield shear and shear viscosity increased with decreasing temperature below 

the lipid melting temperature, showing that the more solid a membrane, the more difficult it was to 

deform in shear. It has been established that, for crystalline materials in general, the yield shear and 

shear viscosity reflect density and mobility of crystal defects in the solid structure. In 1987, this 

team postulated that grain boundaries and/or intra-grain dislocations allowed the shear deformation 

for the 2-molecule thick wax membrane. 

It was expected that because crystalline solids nucleate in liquids upon cooling, the solid lipid 

membrane should form as grains (analogous to panels in a soccer ball; see Figure 34.12) that meet 

at grain boundaries (stitches in the soccer ball). This grain-faceted nanostructure was observed in 

subsequent electron micrograph images of the 100 nm diameter liposome formulation containing 

65 2 

A 

Liquid 

(liquid 

crystalline) 

Waxy lipid membrane 

2 molecules (4 nm) thick 

40 2 A 

40A 

Tc 54A 

Solid 

(gel) 

100 nm in diameter 

100 nm 

FIGURE 34.12 Schematic showing the soccer ball-like structure of the waxy liposome (center). The electron 

micrograph images show solid-phase liposomes, 100 nm in diameter, that are loaded with doxorubicin. The 

nanoscale grains and grain boundaries are evident from the faceted appearance. The phoshpholipid bilayer is 2 

molecules thick (w4 nm), and there is a difference in molecular area and thickness between solid and liquid 

phase lipids (left). (From Ipsen, J. H. and O. G. Mouritsen, Biochim. Biophys. Acta, 944 (2), 121–134, 1988.) 


696 

(a) (b) (c) 

5 μm 

1 μm 

40 μm 

FIGURE 34.13 Solid lipid monolayers on gas microbubbles also show grain structure as revealed by freeze 

fracture electron microscopy (a and b, 5 mm and 1 mm scale). (c) Fluorescence micrograph images show that 

Bodipy lipids included in the lipid formulation at w1 mol% are excluded from the solid grains and forced 

to accumulate at the grain boundaries. (From von Dreele, P. H., Biochemistry, 17 (19), 3939–3943, 1978. 


doxorubicin (Figure 34.12). 149 It was confirmed that this grain structure was evident at the nanoscale 

(100 nm) and for the 2-molecule thick membrane. Also shown in Figure 34.12 is a schematic of 

the solid/liquid interface, illustrating the mismatch in lipid molecular area and thickness between 

solid and liquid states. It is this nanoscale defect that provides the anomalous membrane permeability 

that has been observed and modeled in the transition region 125,126,150 and that is the 

underlying key to the LTSL drug release mechanism. 

Additional evidence for the grain structure came from micromechanical and ultrastructural 

studies of solid lipid monolayers at microgas bubble interfaces (using similar saturated PC 

lipids). This work aimed to create and characterize lipid-stabilized gas-filled ultrasound contrast 

agents. 151 It was observed that the yield shear and shear viscosity were consistent with a polycrystalline 

monolayer, and both freeze fracture electron microscopy and fluorescence microscopy 

showed the grain structure in these micro-lipid shells (Figure 34.13). 

With the grain structure now characterized, the goal of enhancing the permeability at the phase 

transition (compared to pure DPPC and other lipid mixtures) was now the priority. A series of 

micropipette experiments measured and modeled the exchange of molecules that entered the 

bilayer interface (tannic acid) 152 or incorporated into or even crossed the bilayer (peptides 153 

and lysolipids) 142,154 and could then be washed out with bathing media. By measuring the area 

change of a single vesicle and by knowing the areas per molecule of the SOPC * bilayer lipid (65A 2 ) 

and the MOPC † lysolipid (50A 2 ), the mol fraction of lysolipid entering the bilayer from the aqueous 

solution was measured, down to a fraction of a mol%, was measured. Importantly, the desorption of 

lysolipid upon wash-out with lysolipid-free media was also measured, finding that it leaves the 

outer monolayer rapidly (within seconds) 142 . This team then discussed the possibility that the 

membrane permeability to small drugs might be increased if a certain amount of water-soluble 

lysolipids were incorporated in the solid bilayer and then desorbed from the membrane upon 

heating through the transition. The lysolipid MPPC ‡ is structurally and chemically compatible 

with the bilayer lipid DPPC § , but it is quite soluble in water, especially as micelles. It was expected 

that MPPC would leave the bilayer (or form defects) at the T m, making the membrane leaky to 

encapsulated drug. 

Therefore, this team arrived at the new tentative liposome design: a 100 nm liposome 

comprised of a DPPC bilayer (underlying phase transition at 41.98C) with 4 mol% of 

DSPE–PEG (for extended circulation half life) and 10 mol% of the lysolipid MPPC (to enhance 

the permeability), containing doxorubicin loaded by remote pH loading. The key to the design was 

* 

SOPC, Stearolyoleoylphosphatidylcholine. 

† 

MOPC, Monooleoylphosphatidylcholine. 

‡ 

MPPC, Monopalmitoylphosphatidylcholine. 

§ 

DPPC, Dipalmitoylphosphatidylcholine. 




to work with the recognized anomalous permeability at the Tm (an apparent grain boundary effect in 

the melting solid membranes) and to enhance it with a second component (lysolipid). Ultimately, if 

this change in composition could change the membrane property of permeability (when heated), the 

drug would be freely available, and the performance of the temperature-sensitive liposome in vivo 

could be enhanced. 

This lysolipid-based design was tested in 100 nm liposomes, and it characterized their ability to 

become permeable to small molecules, ions, and the drug doxorubicin. By incorporating 10 mol% 

of MPPC or MSPC * into the gel phase bilayers of DPPC liposomes, the permeability rate to 

carboxyfluorescein (CF) 148 and dithionite 115 as well as the rate of release of encapsulated drugs 

like doxorubicin 115,123,155 was greatly enhanced over that of the pure DPPC bilayer at its phase 

transition. For a liposome formulation composed of DPPC:MPPC:DSPE–PEG-2000 † (molar ratio, 

90:10:4), 60% of the entrapped content of CF was released in 1 min at 408C, whereas pure DPPC 

liposomes released less than 20% CF even in 5 min at 408C. Using the dithionite permeability 

assay, 156,157 several lipid systems were compared, providing some very interesting results. In this 

assay, the addition of dithionite quenches the fluorescence and produces a decay in the absorbance 

spectrum of NBD lipids ‡ that are incorporated in the lipid bilayers at low (1 mol%) concentration 

(Figure 34.14). 156 In these experiments, the interaction between dithionite and the NBD-lipid head 

groups was recorded as a decay in absorbance at 465 nm in order to avoid attenuation of fluorescence 

because of repeated exposure to exciting wavelengths. If the bilayer is permeable, the 

absorbance shift will take place on both sides of the bilayer. Otherwise, if the bilayer is impermeable 

(e.g., in its gel state), only the outer monolayer is affected. The change in relative absorbance of 

the outside monolayer upon addition of dithionite matched its relative lipid distribution (i.e., 54% 

outside versus 46% remaining inside), reflecting the high 50 nm radius of curvature of the liposome. 

115 As expected, the permeability for pure DPPC was somewhat enhanced at 40, 42, and 438C 

in the region of its phase transition compared to lower temperatures of 30 and 378C in the gel phase 

region (Figure 34.14). With the inclusion of 10 mol% lysolipid in the bilayers, the absorbance shift 

occurred much faster than for pure DPPC alone, showing that the permeability at the phase 

transition was dramatically enhanced. 115 

Figure 34.15 compares dithionite permeability for several lipid compositions, represented as 

permeability rates (per second). The permeability for DPPC at its Tm (41.98C) was enhanced 

compared to its gel state, but it was surprisingly only slightly higher than that of a liquid phase 

lipid (POPC § ) at 41.98C. 115 At temperatures above 41.98C, both DPPC and POPC exhibited permeability 

rates greater that that of DPPC at its phase transition. In contrast, dramatic increases in the 

phase transition permeability rate were observed when 10 mol% MPPC or MSPC were incorporated 

in the DPPC liposome membranes. The incorporation of these acyl chain-compatible 

lysolipids only shifted the bilayer phase transition temperature slightly from 41.98C for DPPC 

alone to 41.08C with MPPC and 41.48C with MSPC. 115 Therefore, the incorporation of 10 mol% 

lysolipid had achieved the goals of enhancing the transition permeability and slightly lowering the 

transition temperature. In line with the expected influence of the acyl chain, the slightly longer 

MSPC lipid reduced the Tm less than the MPPC. Differential scanning calorimetry (DSC) studies 

showed that the transition region was still very sharp (i.e., not broadened as with the previous 

additions of DSPC and Cholesterol). 148 

A comparison of the DSC data with the permeability rate data confirms that the grain 

boundaries are responsible for the drug release. As shown in Figure 34.16, the permeability rate 

rises on the low temperature shoulder of the excess heat flow curve, i.e., before any significant mass 

* 

MSPC, Monostearoylphosphatidylcholine. 

† 

DSPE–PEG, Distearoylphosphoethanolamine–Polyethyleneglycol 2000. 

‡ 

Lipids conjugated with the fluorescent molecule NBD, N-(7-nitro-2,1,3-benzoxadiazol). 

§ Palmitoyloleoylphosphatidylcholine. 


698 

– 

SO2 – 

SO2 – 

SO2 NBD – 

NBD – 

– 

SO2 – 

SO2 –NBD 

– 

SO2 NBD – 

– 

SO2 – NBD 

– 

SO2 –NBD 

– 

SO2 – 

SO2 – 

SO2 Relative absorbance at 465 nm 

1 

0.9 

0.8 

Dithionite added 

DPPC 30C 

37C 

40C 

0.7 

42C 

0.6 

0.5 

0.4 

0.3 

0.2 

0.1 

0 

43C 

0 2 4 6 8 10 12 14 16 18 20 22 24 26 28 30 

Time (min) 

(a) (b) (c) 

of solid phase lipid material has melted. Also, the maximum permeability rate coincides with the 

midpoint of the transition where it is expected that the grain boundary area is also at its 

maximum. 126 

As for the specific mechanism for this dramatic enhancement of temperature-triggered permeability 

by including 10 mol% lysolipid, it was initially hypothesized that desorption of the watersoluble 

lysolipid would leave defects in the membrane. However, Mills et al. recently showed that 

membrane permeability to dithionite in the transition region was still enhanced following dialysis 

against lysolipid-free buffer (with one exchange of buffer) at temperatures above T m for 48 h 

(Figure 34.17). 115 Under these conditions, with a massive concentration gradient between liposomes 

and dialysate, there was a strong driving force for lysolipid desorption. This study 

demonstrated that, contrary to the original hypothesis, lysolipids did not completely desorb from 

Relative absorbance at 465 nm 

1 

0.9 

0.8 

DPPC:MSPC 

Dithionite added 

30C 

37C 

40C 

0.7 

42C 

0.6 

0.5 

0.4 

0.3 

0.2 

0.1 

0 

43C 

0 2 4 6 8 10 12 14 16 18 20 22 24 26 28 30 

Time (min) 

FIGURE 34.14 The dithionite assay for membrane permeability. (a) Dithionite quenches the fluorescence or, 

in this case, shifts the absorbance for NBD lipids incorporated in the bilayer. The outer monolayer is immediately 

changed, but the inner monolayer is only affected when the membrane becomes permeable to this ion, 

i.e., in the phase transition region. Both compositions show only outer monolayer shifts at temperatures in their 

gel state. Pure DPPC (b) shows the anomalous permeability associated with the transition, but for 

DPPC:MSPC, (c) the change in absorbance occurs much faster. (From Evans, E. and Needham, D., J. Phys. 

Chem., 91 (16), 4219–4228, 1987. With permission.) 

Dithionite ion permeability rate (1/min) 

0.4 

0.35 

0.3 

0.25 

0.2 

0.15 

0.1 

0.05 


DPPC 

DPPC:MPPC (10%) 

POPC 

DPPC:MSPC (10%) 

0 

29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 

Temperature ( o C) 

FIGURE 34.15 Dithionite ion permeability rates for pure and lysolipid-containing TSL membranes in vitro. 

At the phase transition temperature, permeability rates are w10-fold higher for the lysolipid-containing TSLs 

(LTSLs) when compared to the pure DPPC bilayer. (From Mills, J. K. and Needham, D., Biochim. Biophys. 

Acta, 1716 (2), 77–96, 2005. With permission.) 



Heat flow (Endo) 

DPPC:MSPC 

0.00 

30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 


FIGURE 34.16 Comparison between the DSC thermal profile and the dithionite permeability rate for LTSL 

(DPPC:MSPC). This comparison shows that the onset of increased permeability (dashed line) occurs at the low 

temperature shoulder of the excess heat curve (solid line). That is, the content release occurs before a large 

fraction of the lipid has melted. (From Mills, J. K. and Needham, D., Biochim. Biophys. Acta, 1716 (2), 77–96, 


the membrane upon heating to and through the transition; rather, they largely stayed in the 

membrane, apparently forming lysolipid-stabilized defects in the bilayer. 115 Because lysolipids 

(MPPC and MSPC) have a conical molecular shape as a result of a single acyl chain, they prefer 

the micellized state (Figure 34.17b). Therefore, it seems reasonable that they would accumulate at 

Dithionite permeability 

rate constant (1/min) 

(a) 

0.4 

0.35 

0.3 

0.25 

0.2 

0.15 

0.1 

0.05 

DPPC 

DPPC:MPPC10% 

DPPC:MPPC (10%)-Following dialysis 

0 

30 32 34 36 38 

Temperature ( o 40 42 44 46 48 

C) 

shape of lipid 

molecule 

(b) 

Liquid phase region 

(c) 

0.35 

0.30 

0.25 

0.20 

0.15 

0.10 

0.05 

packing of lipid 

molecules 

Permeability rate (1/min) [ ] 

water 

Interface region 

lysolipid-stabilized pore 

Gel phase region 

lipid 

micelle 

lipid 

bilayer 

Gel phase 

DPPC 

Liquid phase 

DPPC 

Lysolipid 

FIGURE 34.17 (a) Dithionite ion permeability as a function of temperature for freshly prepared LTSLs and 

for LTSLs that have been dialysed for 48 h at 458C–488C (liquid state). (b) The maintenance of enhanced 

(although slightly reduced) permeability after dialysis shows that some lysolipid must remain in the 

membrane. Pure DPPC is shown for comparison. (c) Lysolipids normally exist in aqueous solution as micelles 

(left), but they appear to remain present in liposomes at the grain boundaries, possibly stabilizing pores (right). 

(From Mills, J. K. and Needham, D., Biochim. Biophys. Acta, 1716 (2), 77–96, 2005. With permission.) 


700 

melting grain boundaries, forming semispherical porous edges (Figure 34.17c). The resulting 

lysolipid-stabilized pores would form only upon heating to the phase transition region, facilitating 

rapid temperature-sensitive release of contents. Previous micropipet studies by Zhelev and 

Needham on giant lipid vesicles support this hypothesis, suggesting that lysolipid can stabilize 

membrane nanopores large enough to pass glucose, but not sucrose, and that this pore formation 

does not compromise mechanical stability. 158 Therefore, the incorporation of lysolipids (MPPC or 

MSPC) is essential for drug release enhancement for temperature sensitive liposomes. 

For the LTSL formulation, doxorubicin (Dox) was readily loaded into preformed gel phase 

liposomes by remote pH gradient methods (as described in Mayer, Bally, and Cullis and Madden, et 

al. but used now at 358C). 115,159,160 The liposomes were then heated, and Dox release was measured 

as a function of time and temperature. The same triggered permeability and enhanced release 

effects were seen for Dox as with carboxyfluorescein and dithionite. For liposome compositions 

of DPPC:MPPC:DSPE–PEG-2000 (90:10:4) and DPPC:MSPC:DSPE–PEG2000 (90:10:4), 

w100% and w90%, respectively, of the encapsulated Dox was released within 2 min (at 40 and 

41.38C, respectively). 115 This team’s most recent work has improved the experimental conditions in 

order to more closely observe drug release in the first few minutes. 160,161 With the incorporation of 

8 mol% MPPC in the lipid bilayers, 100% of the drug is released in 5 min whereas with 12 mol% 

MPPC, it takes less than 20 s (Figure 34.18). 

The all-important temperature dependence for Dox release was also characterized using 

10 mol% lysolipid as shown in Figure 34.19. At 30, 37, and 398C, very little drug was released. 

At 408C, there was a moderate level of release and at 41.38C, 100% of the drug was released in 4 min. 

If this 4 min time point is used to plot the amount released versus temperature (Figure 34.20), 148 then 

the optimal temperature range for triggered release is 398C–428C—a temperature window that is 

attainable using regional hyperthermia, even for deep-seated tumors such as ovarian and prostate 

tumors. 118 This figure also verifies that, as expected, the stealth liposomes do not release drug in a 

temperature-triggered fashion. Finally, when the Dox release curve is compared to the DSC curve for 

the same liposomes (Figure 34.21), it can be seen that drug release occurs on the low temperature 

shoulder of the phase transition, before a significant mass of lipid is melted. As with the dithionite 

data shown earlier, this suggests a grain boundary mechanism for the drug permeability. 

Importantly, this in vitro data shows that the release time for the lysolipid-containing temperature 

sensitive liposome (LTSL) is likely to be faster than the transit time of the liposome through the 

tumor circulation. Moreover, release can be triggered at temperatures in the mild clinical 

Percent Dox release 

100 

80 

60 

40 

20 

Increasing MSPC 

0 

0 5 10 15 20 25 30 

Time (min) 


0% MSPC 

5% MSPC 

7.4% MSPC 

8.5% MSPC 

9.3% MSPC 

9.7% MSPC 

11% MSPC 

12% MSPC 

13% MSPC 

15% MSPC 

FIGURE 34.18 Drug release as a function of the molar fraction of lysolipid in the bilayer (from 0 mol% to 

15 mol%) at 41.58C. Complete content release occurs within 20 s with 12 mol% MSPC. (From Wright, A., In 

Mechanical Engineering and Materials Science. Duke University, Durham, NC, 2006. With permission.) 



4 10 –9 

100% Release 

3.86 x –9 mol 

3.5 10 –9 

Moles of DOX released 

3 10 –9 

2.5 10 –9 

2 10 –9 

1.5 10 –9 

1 10 –9 

5 10 –10 

0 

0 

41.3 o C 

hyperthermia range. These advantages make the LTSL a much more feasible system for tumor drug 

release than any previous liposome. With this in vitro success, this team was very interested to see 

what effects the LTSL would have on tumor drug accumulation and tumor growth delay in vivo. 

34.3.4 IN VIVO GROWTH DELAY STUDIES 

42 o C 

45 o C 

40 o C 

The efficacy of the Dox–LTSL * was first evaluated in a human squamous cell carcinoma (FaDu) 

model, showing unprecedented local control for this tumor (in two studies, respectively, showing 11 

out of 11 and 6 out of 9 complete regressions, Figure 34.23). 98,123 The FaDu xenografts were grown 

in the flanks of nude athymic mice. All animals received an equivalent single dose of 5 mg/kg i.v. 

doxorubicin (maximally tolerated dose) and one hour of local hyperthermia at 428C. In addition to 

Dox–LTSLCHT, a number of control treatments were tested for comparison. As shown in 

Figure 34.22, with saline control, the FaDu tumors took an average of 10 days to reach the end 

point of 5 times initial tumor volume, representing a fairly fast growing tumor. 98 Administration of 

free Dox showed an average growth time of 12 days, delaying the growth by w2 days.Mild 

hyperthermia (HT) alone, in the absence of any drug, exhibited a growth time of 18 days. When 

free Dox was combined with HT, the growth time was extended by another 4 days to an average of 

22 days. 98 Therefore, these experiments demonstrate the anti-tumor activity of mild hyperthermia 

alone as well as its ability to enhance drug efficacy (as depicted schematically in Figure 34.11). 

In the same FaDu studies, doxorubicin was encapsulated in Doxil-like, non-temperature sensitive 

liposomes (NTSL) as well as in the lysolipid-containing temperature sensitive liposomes 

(LTSL). Figure 34.23a shows the extended growth time of the NTSL in combination with mild 

HT (35 days) with all tumors decreasing rapidly in size but eventually growing back. 123,† Compared 

39 o C 

5 10 15 

Time (min) 

37 o C 

30 o C 

20 25 30 

FIGURE 34.19 Drug release as a function of temperature. For liposomes with 10 mol% MSPC, the rate of 

Dox release increases as a function of temperature with maximum release occurring at the Tm of the lipid 

mixture. (From Needham, D. and Dewhirst, M. W., Adv. Drug Deliv.Rev., 53 (3), 285–305, 2001; Needham, 

D., et al., Cancer Res., 60 (5), 1197–1201, 2000. With permission.) 

* 

Dox–LTSL, doxorubicin-encapsulating lysolipid-containing temperature sensitive liposome. 

† 

The liposome formulations alone, without heat, did not perform much better than the free drug controls, showing that 

reduced accumulation and well-entrapped drug is hardly that efficacious. 


702 

Percent of DOX released 

100 

90 

80 

70 

60 

50 

40 

DOXIL 

DPPC:MSPC(10%) 

30 

20 

10 

0 

30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 



Clinical 

hyperthermia 

range 

FIGURE 34.20 Percent of Dox released within 4 min as a function of temperature for DOXIL-like liposomes 

and LTSL (DPPC:MSPC 10%). For LTSL, triggered release occurred in the clinically attainable temperature 

range of 398C–428C. Over the same temperature range, the Doxil-like liposome (non-temperature-sensitive) 

did not release any drug. (From Needham, D. and Dewhirst, M. W., Adv. Drug Deliv. Rev., 53 (3), 285–305, 

2001; Needham, D., et al., Cancer Res., 60 (5), 1197–1201, 2000. With permission.) 

to NTSL without heat (21 days, not shown), this again proves the synergism of HT and liposomal 

drug delivery through increased perfusion and extravasation as demonstrated in previous 

studies. 106,109,117,162 The tumor regrowth pattern with NTSL indicates that drug was certainly 

delivered, but some cells were able to survive the treatment, perhaps as a result of the relatively 

slow leakage of bioavailable drug. The most dramatic result was observed for the LTSL in combination 

with HT with complete regressions in all 11 animals for up to 60 days (Figure 34.23b). 123 , * 

The importance of rapid drug release was quite apparent. The focus on composition, structure, and 

properties for the liposome design and the testing of in vitro performance of the LTSL translated 

into a significant improvement in tumor response in vivo. 

Following the success of these anti-tumor efficacy studies, additional FaDu experiments quantified 

drug delivery to the tumor and to the actual intracellular molecular target. Total drug 

deposition, tissue distribution, and fraction bound to DNA/RNA were determined for free Dox 

and three different liposome formulations (Doxil-like NTSL † , LTSL, and a more traditional TSL 

with a slightly higher melting point). 98 To do this, two extraction procedures were carried out on 

tumors harvested just one hour after injection. Extraction with chloroform and silver nitrate identified 

all drug in the tumor tissue. Extraction of the tumor tissue with chloroform only identified drug 

that was free, protein-bound, or still trapped inside the liposomes. The difference represented the 

drug that was bound to the polynucleic acids. Dox concentration was measured using high performance 

liquid chromatography (HPLC). The result was that only the LTSL formulation showed 

any doxorubicin bound to DNA and RNA (w50% of the total amount in the tumor, as shown in 

Figure 34.24). The total tumor drug delivery using LTSL plus HT was 25 ng doxorubicin per mg 

tissue (w50 mM). The other liposomes showed 7–8 ng/mg at 428C and w5 ng/mg at 348C. 98 Growth 

time directly correlated with the amount of drug deposited. Tumor tissue sections were also visualized 

for doxorubicin fluorescence to determine the local distribution of the drug in the tumor and 

confirm the relative drug concentrations. 98 The fluorescence intensity was relatively low at 348C 

* 

The second study again showed superior results with six out of nine complete regressions (one tumor disease free at 

40 days) and only two tumors did not respond. 

† 

NTSL, non-temperature sensitive liposomes. 



Heat flow (endo in mW) 

0.2 

0.15 

0.1 

0.05 

0 

0 

30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 


2.5 10 –11 

2 10 –11 

1.5 10 –11 

1 10 –11 

5 10 –12 

FIGURE 34.21 Comparison between the Dox release rate and the DSC curve for DPPC:MSPC (compare 

Figure 34.16). Drug permeability again occurs on the low temperature shoulder of the phase transition before 

significant melting of the lipid mass occurs, indicating a grain boundary mechanism for permeability. 

after free drug or any liposomal treatment group. At 428C, the LTSL showed extensive fluorescence 

(3.09 arbitrary fluorescence units) with the TTSL 3 times less (1.4) and the Doxil-like NTSL 10 times 

less (0.31). These fluorescence images show both unencapsulated and encapsulated drug. For the 

LTSL, the Dox distribution was pervasive throughout these tumor sections (Figure 34.24, right). 

Overall, the LTSL formulation demonstrated increased drug accumulation and triggered drug 

bioavailability, leading to complete tumor regressions after a single dose. 

Relative tumor volume 

10 

8 

6 

4 

2 

0 

10 days 

0 20 40 60 

(a) Days 

(b) 


10 

8 

6 

4 

Control 34°C 

Control 42°C 



2 

0 

18 days 

0 10 20 30 40 50 60 

(c) Days 

(d) 

10 

8 

6 

4 

2 

0 

10 

8 

6 

4 

2 

0 

DOX release rate (Moles/s) 

Free drug 34°C 

12 days 

0 10 20 30 40 50 60 

Days 

Free drug 42°C 

22 days 

0 20 40 60 

Days 

FIGURE 34.22 Tumor growth curves (relative tumor volumes) for flank-implanted FaDu xenografts in nude 

mice. The treatment groups were (a) control at normothermia (348C); (b) free doxorubicin at 348C (maximally 

tolerated dose of 5 mg/kg); (c) 1 h of hyperthermia at 428C; and d) injection of free drug with 1 h HT at 428C. 

The endpoint was time to 5 times initial tumor volume or 60 days. Average growth times are shown for each 

group. Overall, hyperthermia shows independent anti-tumor activity and enhancement of drug activity. (From 

Kong, G., et al., Cancer Res., 60 (24), 6950–6957, 2000. With permission.) 


704 

Relative tumour volume 

(V/Vo ) 

(a) 

Relative tumour volume 

(V/Vo ) 

(b) 

10 

9 

8 

7 

6 

5 

4 

3 

2 

1 

0 

10 

9 

8 

7 

6 

5 

4 

3 

2 

1 

0 

NLTSL 

42°C 

35 days 

5V o 

0 10 20 30 

Days 

40 50 60 

LTSL 42°C 

5V o 

>60 days 

0 10 20 30 40 50 60 

Days 

FIGURE 34.23 Tumor growth curves (relative tumor volumes) for flank-implanted FaDu xenografts in nude 

mice (as in Figure 34.22). All animals received 5 mg/kg doxorubicin and 1 h of local hyperthermia at 428C. 

The treatment groups were (a) non-thermal sensitive Doxil-like liposomes plus HT and (b) Dox–LTSLCHT. 

The Dox–LTSL showed superior anti-tumor efficacy with 11 of 11 complete regressions (up to 60 days). (From 

Needham, D., et al., Cancer Res., 60(5), 1197–1201, 2000. With permission.) 

(a) 

Dox (ng/mg) 

30 

25 

20 

15 

10 

5 

0 

–5 

NTS-Doxil-like 

liposome 

LTS liposome 

TTS liposome 

S34 D34 NT34 TT34 LT34 S42 D42 NT42 TT42 LT42 

Treatment group (34°C or 42°C) 


Amount of Dox bound 

to DNA, RNA 

Amount of Dox still in 

liposomes or bound 

to protein 

FIGURE 34.24 (a) Dox levels in tumor tissue after i.v. infusion as measured by HPLC for each of the 

treatment groups. Hatched bars show the amount still encapsulated or bound to protein, and solid bars show 

the amount bound to DNA/RNA. Only the Dox–LTSL formulation plus HT (428C) shows significant amounts 

of Dox bound to DNA/RNA (w50%). The total tumor [Dox] after LTSL delivery was w50 uM. (right) 

Doxorubicin fluorescence in a representative tumor section immediately prepared after treatment with 1 h 

HT in combination with (b) NTSL, (c) TTSL, or (d) LTSL. Fluorescent intensities are denoted in arbitrary 

units. (From Kong, G., et al., Cancer Res., 60 (24), 6950–6957, 2000. With permission.) 



FIGURE 34.25 A new paradigm for local therapy (compare Figure 34.8). Thermal sensitive liposomes release 

drug RAPIDLY (!20 s) when heated to 428C, resulting in drug release in the blood stream. Therefore, LTSL 

drug delivery does not require liposome extravasation to deliver free drug to the tumor cell. (From Kong, G., 

et al., Cancer Res., 60 (24), 6950–6957, 2000. With permission.) 

Because all three liposomes were 100 nm in size and coated with PEG, they would be expected to 

extravasate to the same extent. However, both the total tumor HPLC results and the histologic 

sections showed additional accumulation of drug (encapsulated plus free) in the tumor with the 

LTSL. This suggested that the greater delivery achieved by LTSL may be due to release of the drug in 

the blood stream of the heated tumor as schematically depicted in Figure 34.25. The Dox concentration 

for even the stealth NTSL (w15 mM) greatly exceeded the 1.8 mMIC50 (media concentration 

required to kill 50% of cells in culture) of free Dox measured in the multi-drug resistant human 

cervical carcinoma line KB-85. 163 However, the limited NTSL efficacy in these studies together with 

the scarcity of complete responses to Doxil in the clinic, suggest that the EPR effect alone has some 

major limitations in drug delivery. As previously mentioned, the in vitro drug release time for LTSL 

is shorter than the transit time of a liposome through tumor microvasculature; therefore, additional 

local delivery through intravascular release would result in free drug diffusion into the tumor tissue. 

Furthermore, endothelial cells lining the tumor microvasculature would also be exposed to free 

doxorubicin. This pointed to a new mechanism for drug delivery, one that does not completely 

rely on the EPR paradigm of liposome extravasation. The possibility of anti-vascular as well as 

anti-neoplastic drug action can now be considered as another aspect of this remarkable paradigm shift. 

34.4 A NEW PARADIGM FOR DRUG DELIVERY TO TUMOR TISSUES 

34.4.1 VASCULAR SHUTDOWN 

A series of preclinical studies in window chamber and flank tumors have established that the 

mechanism of Dox–LTSL action is actually one of vascular shut down. This is due to the unique 

ability of LTSL to intravascularly release Dox as it passes through the blood vessels of the mildly 

heated tumor (Figure 34.25). Utilizing the window chamber for in situ red blood cell velocity 

measurements, Yuan et al. demonstrated that treatment with LTSL plus HT (1 h at 428C) results in 

significantly reduced blood flow immediately after treatment and complete vascular shutdown 

within 24 h in all FaDu tumors (Figure 34.26). 54 Before treatment, blood flow velocity was 

w0.4 mm/s, and at only 6 h after treatment, blood velocity had slowed to w0.002 mm/s. 

Minimal changes in tumor blood flow were seen in all controls, including free drug and empty 

LTSL both with and without HT. 

As shown in Figure 34.27, direct observation of the blood vessels in the tumors before and after 

treatment showed a startling result—not only had blood flow dramatically slowed, but many of the 

tumor microvessels had also disappeared. Within 24 h of LTSL plus the 1 h HT treatment, microvessel 

density in the tumor was reduced from 3.9 mm/mm 2 to 0.9 mm/mm 2 , whereas the control 


706 

(a) 

0.6 

LTSL-HT 

LTSL-DOX-HT 

LTSL-DOX-nHT 

LTSL-nHT 

(b) 

Average RBC velocity 

in tumor vessels (mm/s) 

0.5 

0.4 

0.3 

0.2 

0.1 

0.0 

* 

–0.5 0 6 24 

Time (h) 

* * 

treatments again showed little change (Figure 34.27c). 54 Moreover, these effects were only 

observed in tumor microvessels; heated normal microvessels were spared. Therefore, this represents 

a unique method of functional targeting of therapeutic agents. Measurements of VEGF 

and bFGF expression also showed a 50%–60% reduction in growth factor concentrations in treated 

tumors. 54 Overall, this study provided very strong evidence for the paradigm shift from EPR to 

intravascular release, leading to anti-vascular effects. The EPR effect alone (i.e., extravasation out 

of the blood stream) could not achieve this kind of vascular shut down. The unique ability of the 

lysolipid-containing bilayer to rapidly transition from a solid wax-like membrane into an extremely 

permeable melting state has made possible this rapid intravascular release of drug triggered by the 

application of only mild hyperthermia. This opens up the possibility for drug delivery patterns that 

target both tumor cells and tumor vasculature. 

Average RBC velocity 

in tumor vessels (mm/s) 

0.6 

0.5 

0.4 

0.3 

0.2 

0.1 

0.0 

–0.5 0 6 24 

* 

Time (h) 

LTSL-HT 

LTSL-DOX-HT 

FIGURE 34.26 Red blood cell (RBC) velocity in FaDu tumor microvessels after injection of DiI-labeled 

fluorescent RBCs. (a) For Dox–LTSL plus HT, blood flow was significantly reduced after therapy and was 

essentially zero 24 h later. All controls (empty LTSLGHT, free DoxGHT) showed little effect on blood flow 

velocity. b) For normal vasculature, Dox–LTSLCHT had a much smaller effect. (From Chen, Q, et al., Mol. 

Cancer Ther., With permission.) 

(a) FaDu tumor before treatment 

(b) 

FaDu tumor 24h after treatment 

Microvascular density 

in tumors (mm/mm 2 ) 

(c) 


5 

4 

3 

2 

1 

0 

–0.5 0 6 

Time (h) 

24 

* 

LTSL-HT 

LTSL-DOX-HT 

LTSL-DOX-nHT 

LTSL-nHT 

FIGURE 34.27 Videomicrographs of tumor vasculature showing the tumor blood vessels before (a) and 24 h 

after (b) the LTSL plus 1 h HT treatment. After the treatment, almost all blood vessels disappeared with 

indications of thrombus formation. (c) A plot of tumor microvessel density versus time after the treatment 

comparing LTSLCHT to the same controls as in Figure 34.26. These results point to an anti-vascular 

mechanism found only in tumor vessels. Therefore, the Dox–LTSLCHT therapy appears to be anti-vascular 

as well as anti-neoplastic. (From Chen, Q., et al., Mol. Cancer Ther., 3 (10), 1311–1317, 2004. With permission.) 


**


34.4.2 EXPLORING OTHER TUMOR MODELS 

Dox–LTSL plus mild HT was tested in a number of other tumor models. As with most anti-cancer 

therapies, it was found that different tumors responded to LTSL in different ways. Five tumor lines 

were treated with saline, HT alone, LTSL alone, or LTSL plus HT. 164 Each group received 

equivalent dosing of 6 mg/kg i.v. Dox and/or 1 h HT at 428C. The mouse mammary carcinoma 

4T07 was implanted in normal BalbC mice, and four human tumor xenograft lines were implanted 

in nude athymic mice: SKOV-3 (ovarian carcinoma), HCT116 (colon carcinoma), FaDu (head and 

neck carcinoma), and PC-3 (prostate carcinoma). 164 

All control tumors grew rapidly, reaching 5x tumor volume (5xV) in 10–15 days. HT alone had 

minimal efficacy in delaying tumor growth, as did LTSL alone without heat. The Dox–LTSL plus 

HT groups showed superior tumor growth delay times in all tumor types when compared to all other 

treatment groups (p!0.006), but the degree of prolongation was highly variable from one tumor 

line to the next (Figure 34.28). For Dox–LTSLCHT, the mean tumor growth times (to 5xV) for 

each model were, in order of increasing response, 4T07 (19.8 days)!HCT116 (28.1 days)!FaDu 

(41.8 days)!SKOV3 (51.9 days)!PC3(60 days). 164 This trend correlated to the intrinsic growth 

rates of each untreated control tumor line, i.e., 4T07 and HCT116 were the fastest growing tumors 

and PC was the slowest growing tumor. The fraction of animals remaining disease-free at 60 days 

post treatment was also highly variable among tumor lines. The PC3 tumor line showed the greatest 

response with 15 out of 20 complete regressions. Of note, the FaDu tumor type showed fewer 

complete regressions (1 out of 10) than in previous studies. 98,123 

Potential reasons for these variations in LTSL efficacy were explored by measuring parameters 

describing the tumor microenvironment, tumor cell sensitivity, and overall drug delivery. 164 For 

each tumor type, these parameters were compared to the relative growth delay time and the fraction 

of animals showing complete regression up to 60 days. 

Measurements of pH and pO 2, reflecting the interstitial microenvironment, were performed in 

five mice for each tumor line. The hypothesis that pH might be important was based on the fact that 

tumors produce excess acid. 165 Because Dox is a weak base and is increasingly less hydrophobic 

with decreasing pH, transport of this drug across cell membranes could, in principle, be inhibited by 

low extracellular pH. However, all tumors had the same slightly acidic interstitial pH (6.8–7.0), so 

this could not explain the variations among different tumor lines. The pO2 histograms (100–200 

measurements per tumor) were obtained using the Oxylite probe as described by Braun and 

colleagues. 166 It has been reported that hypoxia can reduce the cyotoxicity of Dox and may 

contribute to different tumor responses in vivo. 167 In this team’s studies, there were significant 

Median Growth Time (Days) 

70 

60 

50 

40 

30 

20 

10 

0 

4T07 HCT116 FaDu SKOV-3 PC-3 

FIGURE 34.28 Median growth time (days to five time tumor volume) for each of five tumor types in response 

to Dox–LTSLCHT. Ranking of median growth time (MGT) reflects the intrinsic tumor growth rates of these 

tumor types. (From Zhao, Y., Yarmolenko, P.S., et al. [In preparation]. With permission.) 


708 

differences in the levels of hypoxia among these tumor lines, ranging from 34 to 64% hypoxic 

fraction. It is expected that the least hypoxic (i.e., most oxygenated) tumors would respond better to 

therapy. This trend was shown in HCT116, SKOV3, and PC3, but FaDu and 4TO7 deviated from 

this pattern (well oxygenated but non-responding). Therefore, the Dox–LTSL efficacy in a given 

tumor line was not completely related to its degree of hypoxia. 164 

The amount of drug delivered to tumors is known to be an important factor in anti-tumor 

efficacy. Therefore, Dox concentration was measured immediately after LTSLCHT treatment in 

eight tumors from each tumor line, using previously published HPLC methods. 98 Drug concentration 

correlated very well for four out of the five tumors, but HCT116 showed the highest drug 

concentration (0.045 mg/mg tumor tissue) along with the poorest response. In addition, cellular 

sensitivity to Dox was determined for each tumor line using the clonogenic assay after 24 h of 

exposure to LTSLCHT. 164 It was found that the in vivo responses of three tumors correlated well 

with in vitro sensitivity to Dox, but two of the most sensitive tumors in vitro (HCT116 and FaDu) 

showed poor response. Altogether, there was not a significant relationship between efficacy and 

either drug concentration or in vitro drug sensitivity. 164 

Overall, the four parameters (extracellular pH, hypoxic fraction, intratumoral drug concentrations, 

and in vitro drug sensitivity) that were examined did not yield a clear pattern to explain the 

differences in Dox–LTSL efficacy among the tumor types. However, searching for positive correlations, 

there were some interesting aspects. PC3 achieved the second highest drug concentration, 

aligning with its high number of complete regressions. In addition, PC3 had the slowest intrinsic 

(control) growth rate and was the most sensitive tumor line in vitro. On the other hand, HCT116 had 

the highest drug concentration, but it yielded only one complete regression (out of 20). The 4T07 

tumor yielded the most consistent explanation for its relative lack of efficacy (no regressions); it had 

the highest IC10 (least sensitive), lowest drug concentration, and fastest intrinsic growth rate 

among all the tumor lines studied. New results from Yuan et al. also show that the relative 

amount of LTSL induced vascular damage is much greater in the FaDu line than in the 4T07 

line (although 4T07 did show some vascular damage, indicating intravascular release of drug from 

liposomes is still an important mechanism). Accordingly, the relative growth delay time (compared 

with control) for FaDu was much higher than for 4T07 (FaDu: 41.8 days/12.1 days Z3.5 vs. 4T07: 

19.8/7.5Z2.6). Therefore, multiple potential explanations for the relative resistance of 4T07 to this 

combined therapy are now available. 

These somewhat inconclusive results demonstrate the complexity of liposomal therapeutics for 

cancer therapy, and this is expected to extend to any other nanotherapeutic agent. For the Dox– 

LTSL, in particular, a new paradigm of drug delivery has now been established, leading to important 

questions regarding mechanisms and methods for a given tumor line and even an individual 

subject. In addition to the parameters already studied, which describe tumor microenvironment, 

drug concentration, and sensitivity of neoplastic cells, it is believed that individual tumor response 

may more critically depend on parameters that specifically affect intravascular drug release and 

vascular shutdown. It is now imperative to evaluate such parameter, including tumor perfusion, 

vascular density, the heatability of vessels, the temperature profile, the location of drug release 

(extravascular versus intravascular), the spatial and temporal distribution of the drug, and the 

relative drug sensitivity of the tumor vascular endothelium (the new target). A comprehensive 

approach is now being taken to evaluate the Dox–LTSL as this is necessary to optimize clinical 

therapy. One challenge that has been faced (that, again, extends to all nanotherapeutics) is the lack 

of tools for real-time evaluation of liposomal drug delivery and drug release at the scale of the 

whole tumor. Though very useful, the window chamber model has been limited to 2D evaluations 

and is not applicable to the clinical setting. This team sought to develop a non-invasive method 

(such as MRI * ) to image liposomal drug concentration distributions in vivo. 

* MRI, magnetic resonance imaging. 




FIGURE 34.29 Median growth time (days to five times tumor volume) for each of five tumor types in response 

to various treatments. The tumor response to Dox–LTSL was greater than all other treatments for all tumor types. 

Ranking of median growth time (MGT) is the same regardless of treatment type for these tumors, reflecting 

intrinsic tumor growth rates. MGTs bound by rectangles are not significantly different from each other 

(Wilcoxon). (From Madden, T. D., et al., Chem. Phys. Lipids, 53 (1), 37–46, 1990. With permission.) 

34.4.3 REGIONAL DEPOSITION OF DRUGS OBSERVED USING MRI 

Through the use of liposomes containing both drug (Dox) and paramagnetic contrast agents 

(manganese), Viglianti et al. showed that LTSL content release and distribution within the tumor 

can be monitored in vivo using T1-weighted MRI. 168 Furthermore, manganese (Mn) and Dox 

concentrations can be calculated using T1 relaxation times and Mn:Dox encapsulation ratios. 168 

This method has been validated in a rat fibrosarcoma model, indicating a concordant linear relationship 

between T1-based [Dox] and invasive [Dox] as measured by HPLC or histologic 

fluorescence. 169 

The potential of this type of imaging for real-time evaluation of therapeutic protocols is 

currently being evaluated, with correlation of drug delivery to outcome on an individual subject 

basis. 170 Figure 34.29 shows representative axial MR images from the rat fibrosarcoma model 

before (a) and after (b) administration of i.v. Mn/Dox–LTSL plus local HT. The flank tumor is 

located in the upper left corner. 169 The catheter at the center of the tumor was used to deliver local 

hyperthermia (via heated water circulation). * Liposomes were injected after the tumor had reached 

steady state temperatures of 408C–428C. The tumor enhancement pattern shows that this protocol 

achieved release of contrast agent (and therefore drug) in the peripheral vasculature,i.e., in the 

feeding vessels as soon as liposomes entered the tumor.. This therapeutic schedule also resulted in 

rapid accumulation of large amounts of drug, likely due to extensive intravascular release (according 

to the new paradigm discussed above). Indeed, the kinetics of drug deposition observed here 

were much faster (!10 min to maximum [Dox]) than the kinetics of liposome extravasation with 

HT (hours to maximum [Dox] using non-thermally sensitive liposomes). 109 This is not surprising, 

since LTSL releases contents within 20 s at 41.38C. 160 Thus, this protocol (inject liposomes during 

HT) targets the rapid drug release specifically to the feeding blood vessels of the tumor, which 

could maximize the anti-vascular effects. Therefore, we hypothesize that the most effective tumor 

response may be achieved with this schedule of therapy. 

Current experiments are evaluating different schedules of LTSL plus HT in terms of drug 

delivery (spatial and temporal tumor distribution) and tumor response (growth delay time) in the 

rat fibrosarcoma model. 170 These results from this kind of study will be important for future clinical 

* 

This is unlikely to be a clinical protocol, but it is a convenient and MR-compatible method to deliver and control local 

heating in animal experiments. 


710 

trials, and preliminary indications are that it may be more beneficial to deliver Dox–LTSL in the 

hyperthermia facility during steady-state heating. Being able to target not only the tumor but also 

the vascular source represents a direct clinical application of this new drug delivery paradigm. 

Furthermore, an imageable temperature-sensitive liposome could be applied in the clinic to monitor 

content release and so facilitate the control of drug distribution and prediction of tumor response for 

a given patient. This new multimodal liposome now provides an anti-neoplastic and anti-vascular 

therapy that can be monitored by MRI in real time. 

34.4.4 CLINICAL TRIALS AND HYPERTHERMIA DEVICES 

Based on the in vitro testing and in vivo pre-clinical success described above, clinical trials are now 

in progress using the Dox–LTSL formulation in combination with local hyperthermia (ThermoDoxe, 

Celsion Corp.). 53 Clinical HT may be delivered to superficial or deep tumors using 

microwave, radiofrequency, or ultrasound devices (Figure 34.30). 118 A phase I trial is currently 

evaluating Dox–LTSL in combination with radiofrequency thermal ablation (RFA) for liver cancer 

(Figure 34.30a). In RFA, the center of the tumor is heated to cytotoxic temperatures greater than 

808C. Whereas this treatment clearly kills the center of the tumor, it is found that tumors often recur 

at the margins after RFA. This occurs because the temperature profile at the margins is only in the 

mild hyperthermia range. Because the Dox–LTSL releases at 398C–428C, it is used in this scenario 

primarily to increase cytotoxicity against these boundary cells. In addition, a phase I/II trial will 

soon test Dox–LTSL in the setting of chest wall recurrence of breast cancer. This trial will use a 

Cool-tip RF 

Thermal zone 

Surgical margin 

Target tissue 

(a) (b) 

(c) (d) 

z (cm) 

0 

−5 

−10 

−15 

−20 

−25 

−20 

0 

20 

x (cm) 

FIGURE 34.30 (See color insert following page 522.) Example of clinical hyper hyperthermia devices. 

(a) The radiofrequency ablation device is percutaneously inserted into the liver tumor under image guidance. 

(b) The superficial microwave applicator may be used for cutaneous or subcutaneous tumors (from http://www. 

ucsfhyperthermia.org/patientinfo/hyperucsf.htm). (c) The Prolieve transurethral microwave applicator was 

developed for benign prostatic hyperplasia but may also be applied for prostate cancer. (d) The RF phased 

array device for breast cancer may be integrated into a table so that the patient may be positioned prone above 

the device. 



−20 

0 

20 

y (cm)


FIGURE 34.31 (See color insert following page 522.) The temperature profile for Celsion’s Prolieve 

transurethral microwave applicator shows that this device can deliver 428C at the periphery of the prostate. 

This enables the device to be used in combination with Dox–LTSL for prostate cancer that typically occurs in 

the periphery. 

superficial microwave applicator to apply mild local HT (BSD500e; Figure 34.30b). Other 

potential clinical applications include prostate cancer (using the Prolievee transurethral microwave 

applicator, Figure 34.30c) and locally advanced breast cancer (using a radiofrequency phased 

array, Figure 34.30d). The Prolievee applicator that was originally designed for benign prostatic 

hyperplasia has demonstrated an ideal temperature profile for the release of LTSL in the periphery 

of the prostate where cancer typically occurs (Figure 34.31). That is, temperatures required for the 

release of drug from circulating liposomes can be attained even at the edges of the prostate by 

heating from the urethra. 


This chapter narrated the development of the lysolipid-containing temperature sensitive liposome 

(LTSL) from bench to bedside in an attempt to highlight many of the formulation, clinical, cellular, 

and in vivo barriers that must be overcome by any nanotechnologically designed cancer therapeutic. 

It has shown that this formulation met an unmet need for bioavailable drug by combining 

previous liposome technologies with a detailed and fundamental knowledge of the materials 

science and engineering (i.e., composition, structure, properties, and processing) of lipid vesicle 

membranes. Historically, the successes of FDA-approved liposomes included reduced toxicity 

compared to free drug administration and a passive tumor accumulation through the EPR effect 

(especially for long-circulating Doxil). However, there were still many barriers to effective drug 

action, and clinical studies have shown limited improvements in tumor response. Reasoning that 

high peaks in bioavailable drug concentration were required for many drugs, this team set out to 

design a liposome with rapid triggered release of contents. In the absence of reliable endogenous 

triggerable chemistry (perhaps excluding phospholipases), 171 of an external trigger (e.g., 

hyperthermia) was utilized with a focus was on local disease rather than global metastases. 

A number of studies contributed to the rationale and design for using hyperthermia to trigger 

liposome release were reviewed. In the window chamber tumor model, the synergistic interactions 

between hyperthermia and liposomal drug delivery were observed, initially for non-thermal sensitive 

liposomes. In addition, many years of fundamental lipid bilayer research revealed important 

relationships among liposome composition, nanostructures (such as crystalline defects in the solid 

state lipid bilayer), and properties (such as enhanced permeability to drugs at the acyl chain melting 

transition). Through the incorporation of lysolipids into the lipid membrane, a dramatic increase in 

permeability at the melting phase transition (418C) was achived, resulting in complete release of 


712 

encapsulated liposomal contents within 20 seconds was reached. The release mechanism appears to 

consist of lysolipid-stabilized pore structures at the interfaces between solid and liquid phase lipids 

at the crystalline grain boundaries. This is supported by the preservation of liposome permeability 

after dialysis and by observations of the material structure and properties of giant liposomes. 

The preclinical studies that have led to clinical testing of the Dox–LTSL in phase I/II trials were 

reviewed. The Dox–LTSL achieved dramatic increases in tumor drug accumulation, up to thirty 

times over free Dox, in mice bearing human squamous cell carcinoma (FaDu) xenografts. This 

resulted in extensive anti-tumor efficacy (17 out of 20 regressions in two initial studies). Mechanistic 

studies in window chamber tumors revealed that Dox–LTSLCHT treatment resulted in 

dramatic reduction in tumor blood flow and complete vascular shutdown for tumors tested. This 

antivascular action was likely due to intravascular release of Dox from the LTSL, representing a 

brand new paradigm for liposomal drug delivery that has previously relied upon extravasation of 

the carrier via the EPR effect. Recent studies in other tumor lines have revealed the complexity and 

heterogeneity of the tumor response to Dox–LTSLCHT. To monitor liposome release and distribution 

in vivo, MRI methods using imageable liposomes have been developed. Ongoing research 

will continue to provide insights as this temperature-sensitive nanotechnology moves forward into 

clinical trials. 

The lessons learned throughout this 15-year effort in triggerable drug delivery are likely to 

extend to other nanotechnological therapeutics. Previous studies have shown that it is possible to 

retain drug in a submicroscopic carrier or attach it to a nanomolecule. These drug delivery systems 

must evade the body’s defenses in order to increase circulation lifetime and tumor uptake. 

However, this chapter has categorically shown that it is also essential to release the drug in free 

form, either in the blood stream of the tumor or in the perivascular space, so that it is bioavailable to 

exert its anti-neoplastic effects as well as important anti-vascular effects. 

REFERENCES 


1. National Library of Medicine. ClinicalTrials.gov [cited Feb 2006]; Available from: http://www. 

clinicaltrials.gov/. 

2. Gabizon, A. A., Stealth liposomes and tumor targeting: One step further in the quest for the magic 

bullet, Clin. Cancer Res., 7(2), 223–225, 2001. 

3. Riethmiller, S., From atoxyl to salvarsan: Searching for the magic bullet, Chemotherapy, 51(5), 

234–242, 2005. 

4. Holland, J. et al., Principles of medical oncology, In Cancer Medicine, Bast, R. et al., Eds., B.C. 

Decker Inc., Hamilton, Ont., 2003. 

5. Willmott, N. and Daly, J., Microspheres and Regional Cancer Therapy, CRC Press Inc., Boca Raton, 

FL, 1994. 

6. Almond, B. A. et al., Efficacy of mitoxantrone-loaded albumin microspheres for intratumoral 

chemotherapy of breast cancer, J. Control. Release, 91(1–2), 147–155, 2003. 

7. Rhines, L. D. et al., Local immunotherapy with interleukin-2 delivered from biodegradable polymer 

microspheres combined with interstitial chemotherapy: A novel treatment for experimental malignant 

glioma, Neurosurgery, 52(4), 872–879, 2003. discussion 879–880. 

8. Allen, T. M. and Cullis, P. R., Drug delivery systems: Entering the mainstream, Science, 303(5665), 

1818–1822, 2004. 

9. Yuan, F. et al., Microvascular permeability and interstitial penetration of sterically stabilized 

(stealth) liposomes in a human tumor xenograft, Cancer Res., 54(13), 3352–3356, 1994. 


microenvironment, Proc. Natl Acad. Sci. U.S.A., 95(8), 4607–4612, 1998. 

11. Torchilin, V. P., Targeted polymeric micelles for delivery of poorly soluble drugs, Cell. Mol. Life 

Sci., 61(19–20), 2549–2559, 2004. 



12. Wang, J., Mongayt, D., and Torchilin, V. P., Polymeric micelles for delivery of poorly soluble drugs: 

Preparation and anticancer activity in vitro of paclitaxel incorporated into mixed micelles based on 

poly(ethylene glycol)–lipid conjugate and positively charged lipids, J. Drug Target., 13(1), 73–80, 

2005. 

13. Kabanov, A. V., Batrakova, E. V., and Alakhov, V. Y., Pluronic block copolymers as novel polymer 

therapeutics for drug and gene delivery, J. Control. Release, 82(2–3), 189–212, 2002. 

14. Rangel-Yagui, C. O., Pessoa, A. Jr., and Tavares, L. C., Micellar solubilization of drugs, J. Pharm. 

Pharm. Sci., 8(2), 147–165, 2005. 

15. Allen, T. M., Liposomal drug formulations. Rationale for development and what we can expect for 

the future, Drugs, 56(5), 747–756, 1998. 

16. Drummond, D. C. et al., Optimizing liposomes for delivery of chemotherapeutic agents to solid 

tumors, Pharmacol. Rev., 51(4), 691–743, 1999. 

17. Discher, B. M. et al., Polymersomes: Tough vesicles made from diblock copolymers, Science, 284 

(5417), 1143–1146, 1999. 

18. Dharap, S. S. et al., Molecular targeting of drug delivery systems to ovarian cancer by BH3 and 

LHRH peptides, J. Control. Release, 91(1–2), 61–73, 2003. 

19. Kopecek, J. et al., Water soluble polymers in tumor targeted delivery, J. Control. Release, 74(1–3), 

147–158, 2001. 

20. Duncan, R. et al., Polymer–drug conjugates, PDEPT and PELT: Basic principles for design and 

transfer from the laboratory to clinic, J. Control. Release, 74(1–3), 135–146, 2001. 

21. Vasey, P. A. et al., Phase I clinical and pharmacokinetic study of PK1 [N-(2-hydroxypropyl)methacrylamide 

copolymer doxorubicin]: First member of a new class of 

chemotherapeutic agents–drug–polymer conjugates. Cancer Research Campaign Phase I/II 

Committee, Clin. Cancer Res., 5(1), 83–94, 1999. 

22. Patri, A. K., Kukowska-Latallo, J. F., and Baker, J. R. Jr., Targeted drug delivery with dendrimers: 

Comparison of the release kinetics of covalently conjugated drug and non-covalent drug inclusion 

complex, Adv. Drug Deliv. Rev., 57(15), 2203–2214, 2005. 

23. Choi, Y. and Baker, J. R. Jr., Targeting cancer cells with DNA-assembled dendrimers: A mix and 

match strategy for cancer, Cell Cycle, 4(5), 669–671, 2005. 

24. Wong, J. Y. et al., A Phase I trial of 90Y-anti-carcinoembryonic antigen chimeric T84.66 radioimmunotherapy 

with 5-fluorouracil in patients with metastatic colorectal cancer, Clin. Cancer Res., 

9(16 Pt 1), 5842–5852, 2003. 

25. Juweid, M. E., Radioimmunotherapy of B-cell non-hodgkin’s lymphoma: From clinical trials to 

clinical practice, J. Nucl. Med., 43(11), 1507–1529, 2002. 

26. Gabizon, A. A. et al., Cardiac safety of pegylated liposomal doxorubicin (Doxil/Caelyx) demonstrated 

by endomyocardial biopsy in patients with advanced malignancies, Cancer Investig., 22(5), 

663–669, 2004. 

27. Berry, G. et al., The use of cardiac biopsy to demonstrate reduced cardiotoxicity in AIDS Kaposi’s 

sarcoma patients treated with pegylated liposomal doxorubicin, Ann. Oncol., 9(7), 711–716, 1998. 

28. Harris, L. et al., Liposome-encapsulated doxorubicin compared with conventional doxorubicin in a 

randomized multicenter trial as first-line therapy of metastatic breast carcinoma, Cancer, 94(1), 

25–36, 2002. 

29. Working, P. K. et al., Reduction of the cardiotoxicity of doxorubicin in rabbits and dogs by encapsulation 

in long-circulating, pegylated liposomes, J. Pharmacol. Exp. Ther., 289(2), 1128–1133, 

1999. 

30. Papahadjopoulos, D. et al., Sterically stabilized liposomes: improvements in pharmacokinetics and 

antitumor therapeutic efficacy, Proc. Natl Acad. Sci. U.S.A., 88(24), 11460–11464, 1991. 


review, J. Control. Release, 65(1–2), 271–284, 2000. 

32. Allen, T. M. et al., Adventures in targeting, J. Liposome Res., 12(1–2), 5–12, 2002. 

33. Bendas, G., Immunoliposomes: A promising approach to targeting cancer therapy, BioDrugs, 15(4), 

215–224, 2001. 

34. Zevalin Injection (Ibritumomab Tiuxetan) Drug Information.[cited; Available from: http://www. 

drugs.com/, 2005. 


714 


35. Witzig, T. E. et al., Randomized controlled trial of yttrium-90-labeled ibritumomab tiuxetan radioimmunotherapy 

versus rituximab immunotherapy for patients with relapsed or refractory low-grade, 

follicular, or transformed B-cell non-Hodgkin’s lymphoma, J. Clin. Oncol., 20(10), 2453–2463, 

2002. 

36. Witzig, T. E. et al., Safety of yttrium-90 ibritumomab tiuxetan radioimmunotherapy for relapsed 

low-grade, follicular, or transformed non-hodgkin’s lymphoma, J. Clin. Oncol., 21(7), 1263–1270, 

2003. 

37. Nademanee, A. et al., A phase 1/2 trial of high-dose yttrium-90-ibritumomab tiuxetan in combination 

with high-dose etoposide and cyclophosphamide followed by autologous stem cell 

transplantation in patients with poor-risk or relapsed non-Hodgkin lymphoma, Blood, 106 (8), 

2896–2902, 2005. 

38. Schilder, R. et al., Follow-up results of a phase II study of ibritumomab tiuxetan radioimmunotherapy 

in patients with relapsed or refractory low-grade, follicular, or transformed B-cell non- 

Hodgkin’s lymphoma and mild thrombocytopenia, Cancer Biother. Radiopharm., 19(4), 478–481, 

2004. 

39. Wiseman, G. A. et al., Ibritumomab tiuxetan radioimmunotherapy for patients with relapsed or 

refractory non-Hodgkin lymphoma and mild thrombocytopenia: A phase II multicenter trial, 

Blood, 99(12), 4336–4342, 2002. 

40. Gibbs, W. J., Drew, R. H., and Perfect, J. R., Liposomal amphotericin B: Clinical experience and 

perspectives, Expert Rev. Anti. Infect. Ther., 3(2), 167–181, 2005. 

41. Astellas Pharma, US., Inc. [cited Feb 2006]; Available from: http://www.astellas.us/press_room/ 

docs/corp_brochure033105.pdf. 

42. Becker, M. J. et al., Enhanced antifungal efficacy in experimental invasive pulmonary aspergillosis 

by combination of AmBisome with Fungizone as assessed by several parameters of antifungal 

response, J. Antimicrob. Chemother., 49, 813–820, 2002. 

43. Schwartz, R. A., Lambert, W. C., Kaposi Sarcoma, Voorhees, A. V. et al., Eds., eMedicine.com, Inc, 

2005. 

44. Armstrong, D. K., Relapsed ovarian cancer: Challenges and management strategies for a chronic 

disease, Oncologist, 7(suppl. 5), 20–28, 2002. 

45. Johnston, S. R. and Gore, M. E., Caelyx: Phase II studies in ovarian cancer, Eur. J. Cancer, 37(suppl. 

9), S8–S14, 2001. 

46. Gordon, A. N. et al., Recurrent epithelial ovarian carcinoma: A randomized phase III study of 

pegylated liposomal doxorubicin versus topotecan, J. Clin. Oncol., 19(14), 3312–3322, 2001. 

47. Berek, J. S. and Bast, R. C., Ovarian cancer, In Cancer Medicine, Bast, R. et al., Eds., B.C. Decker 

Inc., Hamilton, Ont., 2003. 

48. O’Brien, M. E. et al., Reduced cardiotoxicity and comparable efficacy in a phase III trial of pegylated 

liposomal doxorubicin HCl (CAELYX/Doxil) versus conventional doxorubicin for first-line treatment 

of metastatic breast cancer, Ann. Oncol., 15(3), 440–449, 2004. 

49. Muggia, F. M., Clinical efficacy and prospects for use of pegylated liposomal doxorubicin in the 

treatment of ovarian and breast cancers, Drugs, 54(suppl. 4), 22–29, 1997. 

50. Hauck, M. L., Larue, S. M., et al. Phase I trial of doxorubicin-containing low-temperature sensitive 

liposomes in spontaneous canine tumors, Clin. Cancer Res., 12(13) 4004–4010, 2006. 

51. Bangham, A. D., Standish, M. M., and Watkins, J. C., Diffusion of univalent ions across the lamellae 

of swollen phospholipids, J. Mol. Biol., 13(1), 238–252, 1965. 

52. Gregoriadis, G. et al., Drug-carrier potential of liposomes in cancer chemotherapy, Lancet, 1(7870), 

1313–1316, 1974. 

53. Celsion Corp. [cited Feb 2006]; Available from: www.celsion.com/technology/thermodox.cfm. 

54. Chen, Q. et al., Targeting tumor microvessels using doxorubicin encapsulated in a novel thermosensitive 

liposome, Mol. Cancer Ther., 3(10), 1311–1317, 2004. 

55. Allen, T. M. and Smuckler, E. A., Liver pathology accompanying chronic liposome administration in 

mouse, Res. Commun. Chem. Pathol. Pharmacol., 50(2), 281–290, 1985. 

56. Allen, T. M. et al., Chronic liposome administration in mice: Effects on reticuloendothelial function 

and tissue distribution, J. Pharmacol. Exp. Ther., 229(1), 267–275, 1984. 



57. Gallois, L., Fiallo, M., and Garnier-Suillerot, A., Comparison of the interaction of doxorubicin, 

daunorubicin, idarubicin and idarubicinol with large unilamellar vesicles. Circular dichroism 

study, Biochim. Biophys. Acta, 1370(1), 31–40, 1998. 

58. Nichols, J. W. and Deamer, D. W., Catecholamine uptake and concentration by liposomes maintaining 

p/gradients, Biochim. Biophys. Acta, 455(1), 269–271, 1976. 

59. Mayer, L. D., Bally, M. B., and Cullis, P. R., Uptake of adriamycin into large unilamellar vesicles in 

response to a pH gradient, Biochim. Biophys. Acta, 857(1), 123–126, 1986. 

60. Costanzo, L., Board Review Series, Physiology, 2nd ed., Lippincott, Williams and Wilkins, New 

York, 1998. 

61. Truskey, G. A., Yuan, F., and Katz, D. F., Transport Phenomena in Biological Systems, Pearson 

Prentice Hall, Upper Saddle River, NJ, 2004. 

62. Wu, N. Z. et al., Measurement of material extravasation in microvascular networks using fluorescence 

video-microscopy, Microvasc. Res., 46(2), 231–253, 1993. 

63. Takakura, Y., Mahato, R. I., and Hashida, M., Extravasation of macromolecules, Adv. Drug Deliv. 

Rev., 34(1), 93–108, 1998. 

64. Gerlowski, L. E. and Jain, R. K., Effect of hyperthermia on microvascular permeability to macromolecules 

in normal and tumor tissues, Int. J. Microcirc. Clin. Exp., 4(4), 363–372, 1985. 

65. Wu, N. Z. et al., Increased microvascular permeability contributes to preferential accumulation of 

stealth liposomes in tumor tissue, Cancer Res., 53(16), 3765–3770, 1993. 

66. Monsky, W. L. et al., Augmentation of transvascular transport of macromolecules and nanoparticles 

in tumors using vascular endothelial growth factor, Cancer Res., 59(16), 4129–4135, 1999. 

67. DaunoXome Injection Drug Information. [cited; Available from: http://www.drugs.com/], 2005. 

68. Dvorak, H. F. et al., Identification and characterization of the blood vessels of solid tumors that are 

leaky to circulating macromolecules, Am. J. Pathol., 133(1), 95–109, 1988. 

69. Yuan, F., Transvascular drug delivery in solid tumors, Semin. Radiat. Oncol., 8(3), 164–175, 1998. 


cutoff size, Cancer Res., 55(17), 3752–3756, 1995. 


delivery: effect of particle size, Cancer Res., 60(16), 4440–4445, 2000. 

72. Slater, L. M. et al., Cyclosporin A corrects daunorubicin resistance in ehrlich ascites carcinoma, Br. 

J. Cancer, 54(2), 235–238, 1986. 

73. Rivera, E., Liposomal anthracyclines in metastatic breast cancer: Clinical update, Oncologist, 8 

(Suppl 2), 3–9, 2003. 

74. Bradley, A. J. and Devine, D. V., The complement system in liposome clearance: Can complement 

deposition be inhibited?, Adv. Drug Deliv. Rev., 32(1–2), 19–29, 1998. 

75. Allen, T. M., Hansen, C., and Rutledge, J., Liposomes with prolonged circulation times: Factors 

affecting uptake by reticuloendothelial and other tissues, Biochim. Biophys. Acta, 981(1), 27–35, 

1989. 

76. Moghimi, S. M. et al., Non-phagocytic uptake of intravenously injected microspheres in rat spleen: 

Influence of particle size and hydrophilic coating, Biochem. Biophys. Res. Commun., 177 (2), 

861–866, 1991. 

77. Alving, C. R. et al., Therapy of leishmaniasis: Superior efficacies of liposome-encapsulated drugs, 

Proc. Natl Acad. Sci. U.S.A., 75(6), 2959–2963, 1978. 

78. New, R. R. et al., Antileishmanial activity of antimonials entrapped in liposomes, Nature, 272(5648), 

55–56, 1978. 

79. Weldon, J. S. et al., Liposomal chemotherapy in visceral leishmaniasis: An ultrastructural study of an 

intracellular pathway, Z. Parasitenkd., 69(4), 415–424, 1983. 

80. Berman, J. D., U.S. Food and Drug Administration approval of AmBisome (liposomal amphotericin 

B) for treatment of visceral leishmaniasis, Clin. Infect. Dis., 28(1), 49–51, 1999. 

81. Catania, S. et al., Visceral leishmaniasis treated with liposomal amphotericin B, Pediatr. Infect. Dis. 

J., 18(1), 73–74, 1999. 

82. Allen, T. M. et al., Liposomes containing synthetic lipid derivatives of poly(ethylene glycol) show 

prolonged circulation half-lives in vivo, Biochim. Biophys. Acta, 1066(1), 29–36, 1991. 

83. Gabizon, A. and Martin, F., Polyethylene glycol-coated (pegylated) liposomal doxorubicin. 

Rationale for use in solid tumours, Drugs, 54(suppl. 4), 15–21, 1997. 


716 


84. Gabizon, A. et al., In vivo fate of folate-targeted polyethylene–glycol liposomes in tumor-bearing 

mice, Clin. Cancer Res., 9(17), 6551–6559, 2003. 

85. Bedu-Addo, F. K. et al., Effects of polyethyleneglycol chain length and phospholipid acyl chain 

composition on the interaction of polyethyleneglycol–phospholipid conjugates with phospholipid: 

Implications in liposomal drug delivery, Pharm. Res., 13(5), 710–717, 1996. 

86. Torchilin, V. P. et al., Targeted accumulation of polyethylene glycol-coated immunoliposomes in 

infarcted rabbit myocardium, FASEB J., 6(9), 2716–2719, 1992. 

87. Needham, D., McIntosh, T. J., and Lasic, D. D., Repulsive interactions and mechanical stability of 

polymer-grafted lipid membranes, Biochim. Biophys. Acta, 1108(1), 40–48, 1992. 

88. Needham, D., McIntosh, T. J., and Zhelev, D. V., Surface chemistry of the sterically stabilized PEGliposome: 

General principles, In Liposomes: Rational Design, Janoff, A., Ed., Marcel Dekker, Inc., 

New York, pp. 13–62, 1999. 

89. Mori, A. and Huang, L., ”Ninja” liposomes avoiding reticuloendothelial system: Development and 

application, In First Shizuoka Drug Delivery System Symposium. Progress in Drug Delivery Systems, 

University of Shizuoka, Japan, 1992. 

90. Mori, A. et al., Influence of the steric barrier activity of amphipathic poly(ethyleneglycol) and 

ganglioside GM1 on the circulation time of liposomes and on the target binding of immunoliposomes 

in vivo, FEBS Lett., 284(2), 263–266, 1991. 

91. Allen, C. et al., Controlling the physical behavior and biological performance of liposome formulations 

through use of surface grafted poly(ethylene glycol), Biosci. Rep., 22(2), 225–250, 2002. 

92. Papahadjopoulos, D. and Gabizon, A., Liposomes designed to avoid the reticuloendothelial system, 

Prog. Clin. Biol. Res., 343, 85–93, 1990. 

93. Vaage, J. et al., Chemoprevention and therapy of mouse mammary carcinomas with doxorubicin 

encapsulated in sterically stabilized liposomes, Cancer, 73(9), 2366–2371, 1994. 

94. Vaage, J. et al., Therapy of human ovarian carcinoma xenografts using doxorubicin encapsulated in 

sterically stabilized liposomes, Cancer, 72(12), 3671–3675, 1993. 

95. Vaage, J. et al., Tissue distribution and therapeutic effect of intravenous free or encapsulated 

liposomal doxorubicin on human prostate carcinoma xenografts, Cancer, 73(5), 1478–1484, 1994. 

96. Vaage, J. et al., Tumour uptake of doxorubicin in polyethylene glycol-coated liposomes and 

therapeutic effect against a xenografted human pancreatic carcinoma, Br. J. Cancer, 75 (4), 

482–486, 1997. 

97. Doxil Injection Drug Information. [cited; Available from: http://www.drugs.com/], 2005. 

98. Kong, G. et al., Efficacy of liposomes and hyperthermia in a human tumor xenograft model: Importance 

of triggered drug release, Cancer Res., 60(24), 6950–6957, 2000. 

99. Needham, D. and Nunn, R. S., Elastic deformation and failure of lipid bilayer membranes containing 

cholesterol, Biophys. J., 58(4), 997–1009, 1990. 

100. INEX and Enzon announce FDA acceptance of Onco TCS New Drug Application. [cited Feb 2006]; 

Available from: http://www.drugs.com/NDA/onco_tcs_040521.html, 2004. 

101. Krishna, R. et al., Liposomal and nonliposomal drug pharmacokinetics after administration of liposome-encapsulated 

vincristine and their contribution to drug tissue distribution properties, 

J. Pharmacol. Exp. Ther., 298(3), 1206–1212, 2001. 

102. Gelmon, K. A. et al., Phase I study of liposomal vincristine, J. Clin. Oncol., 17(2), 697–705, 1999. 

103. Sarris, A. H. et al., Liposomal vincristine in relapsed non-Hodgkin’s lymphomas: Early results of an 

ongoing phase II trial, Ann. Oncol., 11(1), 69–72, 2000. 

104. Vincristine liposomal—INEX: Lipid-encapsulated vincristine, onco TCS, transmembrane carrier 

system–vincristine, vincacine, vincristine sulfate liposomes for injection, VSLI. Drugs R D, 5(2): 

p. 119–23, 2004. 

105. Huang, Q. et al., Noninvasive visualization of tumors in rodent dorsal skin window chambers, Nat. 

Biotechnol., 17, 1033–1035, 1999. 

106. Matteucci, M. L. et al., Hyperthermia increases accumulation of technetium-99m-labeled liposomes 

in feline sarcomas, Clin. Cancer Res., 6(9), 3748–3755, 2000. 

107. Evans, E. and Needham, D., Physical-properties of surfactant bilayer-membranes–thermal 

transitions, elasticity, rigidity, cohesion, and colloidal interactions, J. Phys. Chem., 91(16), 

4219–4228, 1987. 



108. Kwok, R. and Evans, E., Thermoelasticity of large lecithin bilayer vesicles, Biophys. J., 35(3), 

637–652, 1981. 

109. Kong, G., Braun, R. D., and Dewhirst, M. W., Characterization of the effect of hyperthermia on 

nanoparticle extravasation from tumor vasculature, Cancer Res., 61(7), 3027–3032, 2001. 

110. Wu, N. Z. et al., Simultaneous measurement of liposome extravasation and content release in tumors, 

Microcirculation, 4(1), 83–101, 1997. 

111. Fung, V. W., Chiu, G. N., and Mayer, L. D., Application of purging biotinylated liposomes from 

plasma to elucidate influx and efflux processes associated with accumulation of liposomes in solid 

tumors, Biochim. Biophys. Acta, 1611(1–2), 63–69, 2003. 

112. Satchi-Fainaro, R. et al., PDEPT: Polymer-directed enzyme prodrug therapy. 2. HPMA copolymerbeta-lactamase 

and HPMA copolymer-C-Dox as a model combination, Bioconjug. Chem., 14(4), 

797–804, 2003. 

113. Satchi, R., Connors, T. A., and Duncan, R., PDEPT: Polymer-directed enzyme prodrug therapy. I. 

HPMA copolymer-cathepsin B and PK1 as a model combination, Br. J. Cancer, 85(7), 1070–1076, 

2001. 

114. Needham, D. and Dewhirst, M. W., The development and testing of a new temperature-sensitive 

drug delivery system for the treatment of solid tumors, Adv. Drug Deliv. Rev., 53(3), 285–305, 2001. 

115. Mills, J. K. and Needham, D., Lysolipid incorporation in dipalmitoylphosphatidylcholine bilayer 

membranes enhances the ion permeability and drug release rates at the membrane phase transition, 


116. Rensen, P. C. et al., Determination of the upper size limit for uptake and processing of ligands by the 

asialoglycoprotein receptor on hepatocytes in vitro and in vivo, J. Biol. Chem., 276 (40), 

37577–37584, 2001. 

117. Kong, G. and Dewhirst, M. W., Hyperthermia and liposomes, Int. J. Hyperthermia, 15(5), 345–370, 

1999. 

118. Dewhirst, M. et al., Hyperthermia, In Cancer Medicine, Bast, R. et al., Eds., B.C. Decker Inc., 

Hamilton, Ont., 2003. 

119. Herman, T. S. et al., Effect of heating on lethality due to hyperthermia and selected chemotherapeutic 

drugs, J. Natl Cancer Inst., 68(3), 487–491, 1982. 

120. Estimated incidence, mortality, and site of failure of the most common types of cancer in the United 

States in 2001. Cancer Facts and Figure 2001 in Atlanta: American Cancer Society. 2001. 

121. Suit, H., Assessment of the impact of local control on clinical outcome, Front Radiat. Ther. Oncol., 

29, 17–23, 1996. 

122. Schmidt-Ullrich, R., Local tumor control and survival: Clinical evidence and tumor biologic basis, 

Surg. Oncol. Clin. N. Am., 9(3), 401–414, 2000. See also page vii. 

123. Needham, D. et al., A new temperature-sensitive liposome for use with mild hyperthermia: 

Characterization and testing in a human tumor xenograft model, Cancer Res., 60(5), 1197–1201, 

2000. 

124. Haest, C. W. et al., Fragility of the permeability barrier of Escherichia coli, Biochim. Biophys. Acta, 

288(1), 43–53, 1972. 

125. Papahadjopoulos, D. et al., Phase transitions in phospholipid vesicles. Fluorescence polarization and 

permeability measurements concerning the effect of temperature and cholesterol, Biochim. Biophys. 

Acta, 311(3), 330–348, 1973. 

126. Mouritsen, O. G., Jorgensen, K., and Honger, T., Permeability of lipid bilayers near the phase 

transition, In Permeability and Stability of Lipid Bilayers, Disalvo, E. A. and Simon, S. A., Eds., 

CRC Press, Boca Raton, pp. 137–160, 1995. 

127. Mouritsen, O. G. and Zuckermann, M. J., Model of interfacial melting, Phys. Rev. Lett., 58(4), 

389–392, 1987. 

128. Marsh, D., Watts, A., and Knowles, P. F., Evidence for phase boundary lipid. Permeability of tempocholine 

into dimyristoylphosphatidylcholine vesicles at the phase transition, Biochemistry, 15, 

3570–3578, 1976. 

129. Corvera, E. et al., The permeability and the effect of acyl-chain length for phospholipid bilayers 

containing cholesterol: Theory and experiment, Biochim. Biophys. Acta, 1107, 261–270, 1991. 

130. Nagle, J. F. and Scott, H. L. Jr., Lateral compressibility of lipid mono- and bilayers. Theory of 

membrane permeability, Biochim. Biophys. Acta, 513(2), 236–243, 1978. 


718 


131. Doniach, S., Thermodynamic fluctuations in phospholipid bilayers, J. Chem. Phys., 68, 4912–4916, 

1978. 

132. Jacobson, K. and Papahadjopoulos, D., Phase transitions and phase separations in phospholipid 

membranes induced by changes in temperature, pH, and concentration of bivalent cations, Biochemistry, 

14(1), 152–161, 1975. 

133. Yatvin, M. B. et al., Design of liposomes for enhanced local release of drugs by hyperthermia, 

Science, 202(4374), 1290–1293, 1978. 

134. Magin, R. L. and Niesman, M. R., Temperature-dependent permeability of large unilamellar liposomes, 

Chem. Phys. Lipids, 34(3), 245–256, 1984. 

135. Gaber, M. H., Effect of bovine serum on the phase transition temperature of cholesterol-containing 

liposomes, J. Microencapsul., 15(2), 207–214, 1998. 

136. Ipsen, J. H. and Mouritsen, O. G., Modelling the phase equilibria in two-component membranes of 

phospholipids with different acyl-chain lengths, Biochim. Biophys. Acta, 944(2), 121–134, 1988. 

137. Marsh, D., CRC Handbook of Lipid Bilayers, CRC Press, Boca Raton, 1990. 

138. von Dreele, P. H., Estimation of lateral species separation from phase transitions in nonideal twodimensional 

lipid mixtures, Biochemistry, 17(19), 3939–3943, 1978. 

139. Bloom, M., Evans, E., and Mouritsen, O. G., Physical properties of the fluid lipid-bilayer component 

of cell membranes: A perspective, Q. Rev. Biophys., 24(3), 293–397, 1991. 

140. Zhelev, D. V. and Needham, D., Tension-stabilized pores in giant vesicles: Determination of pore 

size and pore line tension, Biochim. Biophys. Acta, 1147(1), 89–104, 1993. 

141. Olbrich, K., Water permeability and mechanical properties of unsaturated lipid membranes and 

sarcolemma vesicles, In Mechanical Engineering and Materials Science. Duke University: 

Durham, NC, 1997. 

142. Needham, D., Stoicheva, N., and Zhelev, D. V., Exchange of monooleoylphosphatidylcholine as 

monomer and micelle with membranes containing poly(ethylene glycol)-lipid, Biophys. J., 73(5), 

2615–2629, 1997. 

143. Needham, D. and Evans, E., Structure and mechanical properties of giant lipid (DMPC) vesicle 

bilayers from 20 degrees C below to 10 degrees C above the liquid crystal-crystalline phase transition 

at 24 degrees C, Biochemistry, 27(21), 8261–8269, 1988. 

144. Needham, D., McIntosh, T. J., and Evans, E., Thermomechanical and transition properties of dimyristoylphosphatidylcholine/cholesterol 

bilayers, Biochemistry, 27(13), 4668–4673, 1988. 

145. Kim, D. and Needham, D. Lipid bilayers and monolayers: Characterization using micropipet 

manipulation techniques. In Encyclopedia of Surface and Colloid Science, Somasundaran, P., 

Hubbard, A., Et al., Eds., Marcel Dekker, New York, pp. 3057–3086, 2002. 

146. Needham, D. and Zhelev, D., The mechanochemistry of lipid vesicles examined by micropipet 

manipulation techniques, In Vesicles, Rosoff, B., Ed., Macel Dekker, New York, 1996. 

147. Needham, D. and Zhelev, D., Use of micropipet manipulation techniques to measure the properties 

of giant lipid vesicles, In Giant Vesicles, Luisi, P. L., Walde, P., Eds., Wiley, New York, 2000. 

148. Anyarambhatla, G. R. and Needham, D., Enhancement of the phase transition permeability of 

DPPC liposomes by incorporation of MPPC: A new temperature-sensitive liposome for use with 

mild hyperthermia, J. Liposome Res., 9(4), 491–506, 1999. 

149. Ickenstein, L. M. et al., Disc formation in cholesterol-free liposomes during phase transition, 


150. Kruijff, B. D., Demel, R. A., and Deenen, L. L. M. V., The effect of cholesterol and epicholesterol 

incorporation on the permeability and on the phase transition of intact Acholeplasma laidlawii cell 

membranes and derived liposomes, Biochim. Biophys. Acta, 255, 331–347, 1972. 

151. Kim, D. H. et al., Mechanical properties and microstructure of polycrystalline phospholipid monolayer 

shells: Novel solid microparticles, Langmuir, 19(20), 8455–8466, 2003. 

152. Simon, S. A. et al., Increased adhesion between neutral lipid bilayers: Interbilayer bridges formed by 

tannic acid, Biophys. J., 66(6), 1943–1958, 1994. 

153. Zhelev, D. V. et al., Interaction of synthetic HA2 influenza fusion peptide analog with model 

membranes, Biophys. J., 81(1), 285–304, 2001. 

154. Zhelev, D. V., Exchange of monooleoylphosphatidylcholine with single egg phosphatidylcholine 

vesicle membranes, Biophys. J., 71(1), 257–273, 1996. 



155. Wright, A. M., Ickenstein, L. M., and Needham, D. Optimization of a thermosensitive liposome 

formulation for doxorubicin, in preparation. 

156. Langner, M. and Hui, S. W., Dithionite penetration through phospholipid bilayers as a measure of 

defects in lipid molecular packing, Chem. Phys. Lipids, 65(1), 23–30, 1993. 

157. Langner, M. and Hui, S., Effect of free fatty acids on the permeability of 1,2-dimyristoyl-sn-glycero- 

3-phosphocholine bilayer at the main phase transition, Biochim. Biophys. Acta, 1463(2), 439–447, 

2000. 

158. Zhelev, D. V. and Needham, D., Pore formation and pore dynamics in bilayer membranes, ASME. 

Adv. Heat Mass Transfer Biotechnol., 34, 47–49, 1996. 

159. Madden, T. D. et al., The accumulation of drugs within large unilamellar vesicles exhibiting a proton 

gradient: A survey, Chem. Phys. Lipids, 53(1), 37–46, 1990. 

160. Wright, A., Drug loading and release from a thermally sensitive liposome. In Mechanical Engineering 

and Materials Science. Duke University, Durham, NC, 2006. 

161. Wright, A. M., Ickenstein, L. M., and Needham, D., Optimization of a thermosensitive liposome 

formulation for doxorubicin (in preparation Feb 2006). 

162. Gaber, M. H. et al., Thermosensitive liposomes: Extravasation and release of contents in tumor 

microvascular networks, Int. J. Radiat. Oncol. Biol. Phys., 36(5), 1177–1187, 1996. 

163. Gaber, M. H., Modulation of doxorubicin resistance in multidrug-resistance cells by targeted liposomes 

combined with hyperthermia, J. Biochem. Mol. Biol. Biophys., 6(5), 309–314, 2002. 

164. Zhao, Y. et al., Comparative effects of a novel thermosensitive doxorubicin-containing liposome in 

human and mouse tumors. Mol. Cancer Ther. (to be submitted March 2006). 

165. Owen, C. S., Dependence of proton generation on aerobic or anaerobic metabolism and implications 

for tumour pH, Int. J. Hyperthermia, 12(4), 495–499, 1996. 

166. Braun, R. D. et al., Comparison of tumor and normal tissue oxygen tension measurements using 

OxyLite or microelectrodes in rodents, Am. J. Physiol. Heart Circ. Physiol., 280(6), H2533–H2544, 

2001. 

167. Teicher, B. A. et al., Classification of antineoplastic treatments by their differential toxicity toward 

putative oxygenated and hypoxic tumor subpopulations in vivo in the FSaIIC murine fibrosarcoma, 

Cancer Res., 50(11), 3339–3344, 1990. 

168. Viglianti, B. L. et al., In vivo monitoring of tissue pharmacokinetics of liposome/drug using MRI: 

Illustration of targeted delivery, Magn. Reson. Med. 51(6), 1153–62, 2004. 

169. Viglianti, B. L. et al., Chemodosimetry of in-vivo tumor liposomal drug concentration using MRI, J. 

Magn. Reson. Imaging, in press. 

170. Ponce, A. M. et al., Temperature-sensitive liposome release observed by MRI: Drug dose painting 

with hyperthermia impacts anti-tumor effect. Clin. Cancer Res. in press. 

171. Jensen, S. S. et al., Secretory phospholipase A2 as a tumor-specific trigger for targeted delivery of a 

novel class of liposomal prodrug anticancer etherlipids, Mol. Cancer Ther., 3(11), 1451–1458, 2004. 

172. Liposome. 1999 [cited Feb 2006]; Available from: http://www.uic.edu/classes/bios/bios100/ 

lectf03am/liposome.jpg. 

173. O’Driscoll, C., and Duncan, R., Chembytes e-zine: Stealth drugs take off 2000 [cited Feb 2006]; 

Available from: http://www.chemsoc.org/chembytes/ezine/2000/odriscoll_nov00.htm. 

174. Heather Bullen Research Group, Northern Kentucky University. [cited Feb 2006]; Available from: 

http://www.nku.edu/wbullenh1/research.htm. 

175. Kukowska-Latallo, J. F. et al., Nanoparticle targeting of anticancer drug improves therapeutic 

response in animal model of human epithelial cancer, Cancer Res., 65(12), 5317–5324, 2005. 

176. Klement, A., TARA News, High-Tech. Behandlung bei malignem Lymphom: Zevalin. 2004 [cited; 

Available from: http://www.oeaz.at/zeitung/3aktuell/2004/17/serie/serie17_2004tara.html. 

177. Hyperthermia at University of California—San Francisco. [cited Feb 2006.]; Available from: http:// 

www.ucsfhyperthermia.org/patientinfo/hyperucsf.htm. 


35 

CONTENTS 

Nanoemulsion Formulations for 

Tumor-Targeted Delivery 

Sandip B. Tiwari and Mansoor M. Amiji 

35.1 Introduction ....................................................................................................................... 723 

35.2 Composition of Nanoemulsions ........................................................................................ 725 

35.3 Preparation of Nanoemulsions .......................................................................................... 726 

35.4 Characterization of Nanoemulsion Systems ..................................................................... 727 

35.4.1 Particle Size......................................................................................................... 727 

35.4.2 Surface Characteristics ........................................................................................ 728 

35.5 Fate of Nanoemulsions In Vivo ........................................................................................ 729 

35.5.1 Effect of Particle Size on Disposition and Clearance ........................................ 729 

35.5.2 Effect of Surface Characteristics on Disposition and Clearance ....................... 729 

35.6 Applications of Nanoemulsions in Cancer Therapy......................................................... 730 

35.6.1 Improved Delivery of Lipophilic Drugs ............................................................. 730 

35.6.2 Positively Charged Anoemulsions in Cancer Therapy....................................... 731 

35.6.3 Photodynamic Therapy of Cancer ...................................................................... 732 

35.6.4 Neutron Capture Therapy of Cancer .................................................................. 733 

35.6.5 Perfluorochemical Nanoemulsions in Cancer Therapy ...................................... 734 


References .................................................................................................................................... 736 


An emulsion is generally described as a heterogonous system composed of two immiscible liquids: 

one dispersed uniformly as fine droplets throughout the other. The majority of conventional emulsions 

in pharmaceutical use have dispersed particles ranging in diameter from 0.1 to 100 mm. As 

with other dispersed systems such as suspensions, emulsions are thermodynamically unstable as a 

result of the excess free energy associated with the surface of the droplets in the internal phase. The 

dispersed droplets strive to come together and reduce the surface area. In addition to this flocculation 

effect, the dispersed particles can coalesce, or fuse, and this can result in the eventual 

destruction (phase separation) of the emulsion. To minimize this effect, a third component, the 

emulsifying agent or surface active agent (surfactant), is added to the system to improve the 

stability. The liquid present as small globules is called the dispersed or internal phase, whereas 

the liquid where the droplets are dispersed is called the dispersion, continuous, or external phase. In 

general, the immiscible liquids are described as oil and water phases. The oil phase may be any 

hydrophobic non-polar liquid, and the water phase may be any hydrophilic polar liquid. 


723

724 

The term macroemulsion is sometimes employed to distinguish the ordinary emulsions defined 

above from microemulsions, nanoemulsions, and micellar systems. Whereas nanoemulsions are 

two-phase systems where the dispersed phase droplet size has been made in the nanometer size 

range, the microemulsions, and micellar systems are single-phase systems. 1–3 All these systems are a 

result of interfacial phenomenon brought out by surface active agents. Therefore, it is essential to 

understand the molecular mechanisms responsible for the creation of these systems and distinctions 

between them. Surface active compounds contain hydrophilic and hydrophobic moieties in the same 

molecule. When these surfactants are added to the water phase or the oil phase, one portion of the 

molecule is always incompatible with the other solvent molecules. Surfactants overcome this incompatibility 

by adsorption at interfaces (air–water, oil–water, or solid–water) and/or by formation of 

aggregates called micelles. Micelles are colloidal aggregates that form above a specific concentration, 

the so-called critical micelle concentration (CMC). When the solvent is water, a micellar 

aggregate exhibit the structure as shown in Figure 35.1a where the surfactant molecules are oriented 

in such a way that the hydrophilic groups are in contact with the water molecules, and the hydrophobic 

tails are located in the micellar core away from the aqueous environment. In this case, the core 

of a micelle resembles a tiny pool of liquid hydrocarbon where compounds that are poorly soluble in 

water but soluble in non-polar solvents can be solubilized. When the solvent is oil phase, inverse (or 

reverse) micelles are formed where the surfactant’s hydrophilic groups are located inside 

(Figure 35.1b). In this case, the compounds that are soluble in water phase can be solubilized in 

the micellar core. The process of solubilization takes place well above the CMC, and because of 

solubilization, micelles become swollen and may attain the size of a small droplet (w100 nm). 

Consider, for example, that water is solvent, the surfactant concentration is well above CMC, and 

the oil phase is also present; the micelles would solubilize large amounts of the oil phase and become 

swollen until they start interacting through a phenomenon called percolation. Such packed swollen 

micelle structures that could solubilize large amounts of both oil- and water-soluble compounds have 

been called microemulsions because they were first thought to be extremely small droplet emulsions. 

4 This definition of microemulsions is still controversial with some scientists referring 

microemulsions in similar context of emulsions systems (two phase systems, e.g., nanoemulsions) 

while others referring to it as a single phase system. 1,4,5 There is, however, general agreement that 

microemulsions are optically isotropic, thermodynamically stable systems and are not amenable to 

dilution with the external phase as normal emulsions. Application of microemulsions is usually 

limited to dermal and peroral application because of their high surfactant concentration that tends 

to prove toxic through other routes of administration. Also, they exist in narrow regions of the phase 

diagrams and, as such, they are very sensitive to quantitative changes in formulation. A notable 

a. Micelle structure 

(Dispersion medium is water phase) 

Polar head group 

Non-polar tail group 

Emulsier or surfactant molecule 

FIGURE 35.1 (a) Micelles and (b) inverse micelles. 



b. Inverse micelle 

(Dispersion medium is oil 

phase)

Nanoemulsion Formulations for Tumor-Targeted Delivery 725 

example of the microemulsion system in pharmaceutical application is Sandimmune Optorale and 

Neorale pre-concentrate that contains cyclosporin-A as the active moiety indicated for immune 

suppression in organ transplantation. 6 

Emulsion formulations can be used orally, parenterally, ophthalmically, or topically, and they 

can be used by respiratory route. 4,7 Commercial formulations of nanoemulsions are presented in 

Table 35.1. Major application areas of nanoemulsions are in the field of total parenteral nutrition 

and as carriers for various drugs, anesthetics, and image contrast-enhancement agents. In this 

chapter, the discussion on application of nanoemulsion systems is restricted to cancer therapy. 

Two types of emulsion systems are possible depending on if the dispersed droplets are oil or water 

(i.e., oil-in-water (O/W) or water-in-oil (W/O) emulsion). O/W nanoemulsion systems are of 

particular pharmaceutical importance from a parenteral drug delivery point of view. W/O emulsions 

are also parenterally used; however, their applications are limited for obtaining the prolonged 

release of the water soluble compounds on intramuscular application. Multiple emulsion or double 

emulsion systems are also possible where preformed emulsion is re-emulsified with either oil or 

water. For example, multiple emulsions of oil-in-water-in-oil (O/W/O) are W/O emulsions where 

the water droplets contain dispersed oil globules. Similarly, water-in-oil-in-water (W/O/W) 

multiple emulsions are those where the internal and external water phases are separated by an 

oil phase (Figure 35.2). 

35.2 COMPOSITION OF NANOEMULSIONS 

Nanoemulsions contain oil phases, surfactants or emulsifiers, active pharmaceutical ingredients 

(drugs or diagnostic agents), and additives (Table 35.2). 5,8–10 The oil phases are mainly natural or 

synthetic lipids, fatty acids, oils such as medium or long chain triglycerides, or perflurochemicals. 

Many oils, in particular, those of vegetable origin, are liable to auto-oxidation, and their use in 

pharmaceutical formulations requires the addition of an antioxidant. The most widely used oils for 

parenteral applications are purified soybean, corn, castor, peanut, cottonseed, sesame, and safflower 

oils. Table 35.1 lists some of the oils that can be used for formulating nanoemulsions. Squalene has 

been reported to be the choice of oil for formulating stable nanoemulsions with smallest droplet 

size. 11 Squalene, biocompatible oil, is a linear hydrocarbon precursor of cholesterol found in many 

tissues, notably the livers of sharks (Squalus) and other fishes. 12 Squalane, a derivative of squalene, 

is prepared by hydrogenation of squalene and is fully saturated that means that it is not subject to 

auto-oxidation, an important issue from a stability point of view. Purified mineral oil is used in 

some W/O emulsion preparations. 10 Emulsified perflurochemicals are considered acceptable for 

TABLE 35.1 

Commercial Nanoemulsion (Sub-Micron Emulsion) Formulations 

Drug Brand Manufacturer Therapeutic Indication 

Propofol Diprivan Astra Zeneca Anesthetic 

Dexamethasone 

Palmitate 

Limethason Mitsubishi Pharmaceutical, 

Japan 

Steroid 

Alprostadil Liple Mitsubishi Pharmaceutical, 

Japan 

Vasodilator platelet inhibitor 

Flurbiprofen axetil Ropion Kaken Pharmaceuticals, 

Japan 

Nonsteroidal analgesic 

Vitamins A, D, E, K Vitalipid Fresenius Kabi, Europe Parenteral nutrition 


726 

W 

Oil 

Water-in oil 

(W/O emulsion) 

Oil 

Oil-in water 

(O/W emulsion) 

intravenous application, and many formulations involving the use of perflurochemicals have 

reached the marketplace. 13 

The most widely used oils for oral emulsions are various fixed oils of vegetable origin, (corn, 

cotton seed, or peanut oil), fish liver oil, and mineral oil. The most commonly used emulsifiers for 

parenteral application include either natural lecithins derived from plant or animal sources [e.g., 

egg yolk or soy phosphatidylcholine (PC)] or modified lecithins (e.g., diethylenetriaminepentaacetic 

acid (DTPA) conjugated phosphatidylethanolamine (PE) or poly(ethylene glycol) (PEG)modified 

lecithin derivatives). The poly(ethylene oxide) (PEO)-containing block copolymers, in 

particular Pluronics w or poloxamers, Tetronic w or poloxamines, polysorbates, PEG-conjugated 

castor oil derivatives (Cremophore EL w ), polyglycolized as well as glycerides, stearylamine 

(SA), 1, 2-dioleoyl-3-trimethylammonium-propane, N-(1, 2-Dioleyl-dihydroxypropyl)-N-(dihydroxy-propyl) 

(DOTAP) or other positively charged lipids are also used as emulsifiers or 

co-emulsifiers in the nanoemulsion formulations. Other pharmaceutical additives such as tonicity 

modifiers, pH adjustment agents, antioxidants, sweeteners, flavors, and preservatives may also be 

included in the final formulation, depending on the need. 

35.3 PREPARATION OF NANOEMULSIONS 

W 2 

Water-in-oil-in-water 

(W 1 /O/W 2 emulsion) 

W 

Water 

Oil 1 

O 2 

Oil 

Oil-in-water-in-oil 

(O 1 /W/O 2 emulsion) 


Nanoemulsion/ miniemulsion/ 

sub-micron emulsion 

(Type: O/W Nanoemulsion) 

As previously discussed, because nanoemulsions are dispersions of fine droplets (in nanometer 

range) in a dispersion medium, they are thermodynamically unstable and require considerable 

mechanical energy for their preparation. 14–16 The mechanical energy can be supplied in the form 

of high pressure homogenizer (e.g., Avestin w homogenizer, Avestin Inc., Ottawa, Canada), Microfluidizer 

w (MFIC Corporation, Newton, MA, U.S.A.), or an ultrasonic generator (e.g., VC-505, 

Sonics, and Materials Inc., CT, U.S.A.). Stable preparation of nanoemulsions involves the 

following steps. First, the active hydrophobic pharmaceutical agent is dissolved in the oil phase 

and mixed with the water phase, and the mixture is then emulsified with a high shear homogenizer 

for a short period. The coarse oil-in-water emulsion formed is then further homogenized using a 

high pressure homogenizer or Microfludizer w . The emulsifier (or surfactant) is either mixed with 

the water phase or the oil phase. Incorporation of the surfactant in the oil phase has added advantage 

W 

Micro emulsion 

(Single phase systems) 

FIGURE 35.2 The different types of colloidal systems. (Adapted from Salager, J. L., Pharmaceutical Emulsions 

and Suspensions, Marcel Dekker, New York, pp. 19–72, 2000.) 



TABLE 35.2 

Components of Nanoemulsion Formulations 

as it may aid in enhancing solubility of the lipophillic compound in the oil phase. The size of 

nanoemulsion droplets can be manipulated by using appropriate pressure settings and number of 

cycle runs. Nanoemulsions can also be formulated using direct sonication of the oil phase with 

water phase in the presence of emulsifying agent. In this case, the intensity, mode, and duration of 

sonication determine the final droplet size. It should, however, be noted that the size of nanoemulsion 

droplets is also dependant on the nature of oil (e.g., oil composition and viscosity) and the 

emulsifying phase (e.g., type of emulsifier and its concentration). 

35.4 CHARACTERIZATION OF NANOEMULSION SYSTEMS 

Nanoemulsion systems are routinely characterized for particle size and surface properties (surface 

charge and morphology). 

35.4.1 PARTICLE SIZE 

Oils Emulsifiers 

Castor oil Natural lecithins from plant or animal sources 

Coconut oil (phospholipids) 

Corn oil PEG- phospholipids 

Cottonseed oil Poloxamers (e.g. F68) 

Evening primrose oil Polysorbates 

Fish oil Polyoxyethylene castor oil derivatives 

Jojoba oil Polyglycolized glycerides 

Lard oil Stearlyamine 

Linseed oil Oleylamine 

Mineral oil 

Olive oil Additives 

Peanut oil Antioxidants (a-tocopherol, ascorbic acid) 

PEG-vegetable oil Tonicity modifiers (glycerol, sorbitol, xylitol) 

Perflurochemicals pH adjustment agents (NaOH or HCl) 

Pine nut oil Preservatives 

Safflower oil 

Sesame oil 

Soybean oil 

Squalene 

Squalane 

Sunflower oil 

Wheatgerm oil 

Source: From Eccleston, G., Encyclopedia of Pharmaceutical Technology, Marcel Dekker, New 

York, pp. 137–188, 1992; Chung, H., Kim, T. W., Kwon, M., Kwon, I. C., and Jeong, S. C., J. 

Control Release, 71, 339–350, 2001; Allison, A. D., Methods, 19, 87–93, 1999; Klang, S. and 

Benita, S., Submicron Emulsions in Drug Targeting and Delivery, Academic Publishers, 

Harwood, pp. 119–152, 1998; Lundberg, B. B., J. Pharm. Pharmacol., 49(1), 16–21, 1997; 

Campbell, S. E., Kuo, C. J., Hebert, B., Rakusan, K., Marshall, H. W., and Faithfull, N. S., 

Can. J. Cardiol., 7(5), 234–244, 1991. 

The size of nanoemulsion droplets determines their behavior both in vitro and in vivo. Particle size 

of nanoemulsion droplets can be measured using an ensemble (e.g., spectroscopic methods such as 


728 

light scattering), counting (e.g., microscopy such as freeze fracture electron microscopy) or separation 

method (e.g., analytical ultracentrifugation). 17,18 Each method has its own advantages and 

disadvantages. For example, when the particle size measurement of standard nanoemulsions (made 

with pine nut oil and egg PC) was performed by this laboratory using an ensemble and counting 

method, the results indicated some discrepancy. The average particle size of the nanoemulsion 

droplets was 110G13 nm by dynamic light scattering method using ZetaPALS, 90Plus (Brookhaven 

Instruments Corporation, Holtsville, NY, U.S.A.). However, the freeze fracture electron 

microscopy images indicated the size of the droplets in the range of 25–40 nm with some particle 

aggregates in the size range w100–150 nm (Figure 35.3). It is recommended that, in order to obtain 

better information on particle size distribution, it is essential that particle size analysis be performed 

by more than one method. The particle size of the commercial nanoemulsions intended for total 

parenteral nutrition is reported to be in the range of 100–400 nm. 19 

35.4.2 SURFACE CHARACTERISTICS 


Similar to particle size, surface charge of the nanoemulsion droplets has marked effect on the 

stability of the emulsion system and the droplets’ in vivo disposition and clearance. Conventionally, 

the surface charge on the emulsion droplets has been expressed in terms of zeta potential (z). As the 

emulsion droplets are a result of interfacial phenomenon brought out by surface active agents, their 

zeta potential is dependent on the extent of ionization of these surface active agents. According to 

DLVO electrostatic theory, the stability of the colloid is a balance between the attractive van der 

Waals’ forces and the electrical repulsion because of the net surface charge. 20 If the zeta potential 

FIGURE 35.3 Freeze fracture electron microscopy image of paclitaxel nanoemulsions formulated using 20% 

oil phase (pine nut oil) and egg phosphatidylcholine as surfactant and deoxycholic acid as co-surfactant. The 

inset shows microscopy image at higher magnification. Nanoemulsion droplets were in the size range of 

25–40 nm with some particle aggregates in the size range of w100–150 nm. Scale bar in both figures 

represents 100 nm. 



falls below a certain level, the emulsion droplets will aggregate as a result of the attractive forces. 

Conversely, a high zeta potential (either positive or negative), typically more than 30 mV, maintains 

a stable system. 8,15,19 The zeta potential of the nanoemulsion droplets is routinely measured 

using a Zetasizer (Malvern Instruments, U.K.) or the ZetaPlus instrument (Brookhaven Instruments 

Corporation, Holtsville, NY, U.S.A.). The zeta potential of the commercial nanoemulsions used for 

total parenteral nutrition is reported to be around K40 to K50 mV at pH 7. 19 The net negative 

surface charge is thought to be attributed to the anionic components of the emulsions, mainly the 

phospholipids. 21 One can impart positive charge to the nanoemulsions as well using cationic lipids 

[e.g., SA, 22–24 oleylamine, 25 2,3-dioleoyloxypropyl-1-trimethylammonium bromide, DOTMA, 26 

dimethylaminoethane carbamoyl cholesterol (DC-cholesterol)], 27 polymers (e.g., chitosan), 28,29 

and cationic surfactants (e.g., cetyltrimethylammonium bromide). 30 

35.5 FATE OF NANOEMULSIONS IN VIVO 

Similar to liposomal systems, the oil droplets of a conventional O/W nanoemulsion are rapidly 

cleared after intravenous injection by the reticuloendothelial system (RES) organs such as the liver, 

spleen, and bone marrow. Alternatively, the emulsion droplets would bind to the apolipoproteins 

from the blood stream, followed by their metabolism similar to chylomicrons.The clearance rate of 

emulsion droplets from the systemic circulation is dependant on the physicochemical properties of 

the emulsion system (size and surface properties in particular). By appropriate control of these 

physicochemical properties, one can manipulate the in vivo behavior of the nanoemulsion systems. 

35.5.1 EFFECT OF PARTICLE SIZE ON DISPOSITION AND CLEARANCE 

Small-sized emulsion droplets are cleared slower than larger emulsion droplets. It is reported that 

small-sized emulsions (100–110 nm) were successful in enhancing the plasma concentration of 

rhizoxin although the drug was also distributed to the peripheral tissues. 31 On the other hand, 

emulsions with droplet size O200 nm prevented the drug from entering the bone marrow, small 

intestine, and other organs not associated with the RES where toxicities of many cytotoxic 

compounds are expected. Size of nanoemulsion systems can be controlled by processing conditions 

(e.g., number of homogenization cycles or duration and intensity of sonication) and also by composition 

of the emulsifier system. Lundberg has reported that it is possible to control the particle size 

of the nanoemulsion system using appropriate co-surfactant, e.g., use of 0.12 wt% of polysorbate 

80 in castor oil: egg PC (1:0.4 wt%) emulsion system was found to reduce the size of the droplets 

from 50 to 100 nm without using any co-surfactant. 32 The use of Pluronic w F-108 as one of the 

co-surfactants also decreased the size of the emulsion droplets. Particle size is also dependant on the 

type of the oil phase used for the preparation of the nanoemulsion systems. For example, it was 

reported that when squalene was used as the oil phase, smallest nanoemulsion droplets (190 nm) 

were observed. 11 Castor, safflower, and lard oils yielded emulsion droplets in the size range of 270– 

285 nm, whereas linseed oil yielded droplets with size of around 355 nm. 

35.5.2 EFFECT OF SURFACE CHARACTERISTICS ON DISPOSITION AND CLEARANCE 

As previously discussed, emulsion droplets are rapidly taken up by the cells of the RES. The use of 

PEG-modified phospholipids as emulsifies were shown to reduce the uptake by the cells of the RES 

and to improve the blood circulation times. For example, it was reported that inclusion of PEG 

derivatives such as Tween w 80 or PEG–PE into the emulsions composed of castor oil and PC 

decreases RES uptake and improves systemic circulation time. 33 Hodoshima et al. 34 reported that 

lipid nanoparticles of lipophillic prodrug 4 0 -O-tetrahydropyranyl doxorubicin with synthetic PEG 

(1,900)-lipid derivative and PC enhanced the circulation half of the drug to 4.3 h after intravenous 

administration in BALB/c mice. A similar approach has been used to improve tumor delivery of 


730 

rhizoxin. 31,35 When lecithin was replaced with hydrophilic poloxamer 338 in the emulsion formulation, 

it was possible to avoid the normal disposition of droplets to the liver and spleen. In fact, one 

study reported that one could totally eliminate the liver and spleen uptake by using a nonionic 

co-emulsifier, the block copolymer poloxamine (P 908). 36 Emulsions prepared with Pluronics w as 

co-emulsifiers also enhanced the circulation time of the formulation in vivo. The nature of lipid 

component also alters the biodistribution profile of emulsion systems. For example, it was reported 

that cholesterol-coated emulsions were rapidly removed from circulation compared to those coated 

with sphingomyelin. 37 The surface charge of the nanoemulsion droplets also affects the clearance 

rate in vivo. Negatively charged nanoemulsions are cleared faster and showed high RES organ 

uptake than neutral or positively charged nanoemulsions. 8 On the other hand, positively charged 

nanoemulsions showed an initial accumulation in the lungs followed by redistribution to the liver 

and spleen. 15 This natural fate of positively charged nanoemulsions might be important in developing 

a suitable system for delivery of drugs targeted specifically to the lungs in case of transplant 

rejection therapy, lung infections, or lung cancer. In brief, the surface modification of nanoemulsion 

systems offer an opportunity for prolonging the blood circulation time of the drug, enhancing the 

possibility of interaction with target cells of interest. This is, in particular, important for delivery of 

anti-cancer drugs as prolonged exposure of the drug is critical in the treatment of multi-drugresistant 

tumors. 

35.6 APPLICATIONS OF NANOEMULSIONS IN CANCER THERAPY 

35.6.1 IMPROVED DELIVERY OF LIPOPHILIC DRUGS 


The advantages of formulating various lipophilic anti-cancer drugs in submicron O/W emulsion are 

obvious. The oil phase of the emulsion systems can act as a solubilizer for the lipophilic compound. 

Therefore, solubility of lipophilic drugs can be significantly enhanced in an emulsion system, 

leading to smaller administration volumes compared to an aqueous solution. In addition, because 

lipophilic drugs are incorporated within the innermost oil phase, they are sequestered from direct 

contact with body fluids and tissues. Lipid emulsions can minimize the pain associated with 

intravenously administered drugs by exposing the tissues to lower concentrations of the drug or 

by avoiding a tissue irritating vehicle. This has been demonstrated with propofol, 38 diazepam, 39 

methohexital, 40 clarithromycin, 41 etomidate, 42 and a novel cytotoxic agent. 43 Also, because these 

systems are similar to chylomicrons, they are well-tolerated and have low incidences of side effects 

compared to systems based on organic solvents, pH adjustments, and surface active agents (e.g., 

Cremophor) because there is less chance of drug precipitation upon administration. If the drug is 

susceptible to hydrolysis or oxidation, it is protected by the non-aqueous environment. Furthermore, 

incorporation of anti-cancer drugs in submicron emulsions (with droplet size of 50–200 nm) 

with long circulation properties are expected to enhance the tumor accumulation of the drug by 

passive targeting through the enhanced permeability and retention effect. 44,45 It possible to enhance 

the tumor accumulation of nanoemulsions with appropriate modification of size or surface functionalization 

as previously discussed. Oil-in-water submicron emulsions appear to be a viable 

alternative for the intravenous administration of various lipophilic cytotoxic drugs. Several 

groups of researchers have reported the submicron emulsion formulations of anti-cancer drugs 

for improved efficacy and/or reduced toxicity. 

Paclitaxel, a highly potent anti-cancer agent initially extracted from the bark of the Pacific yew, 

was entrapped in lipid emulsion droplets with triolein as oil core and dipalmitoylphosphatidylcholine 

(DPPC) as the principal emulsifier. The emulsion was stabilized with polysorbate 80 and PEGdipalmitoyl 

PE. The results of in vitro anti-proliferative activity of paclitaxel against T47D cells 

was retained by the drug emulsion with IC50 (drug concentration effective in reducing proliferation 

by 50%) value of 7 nM as compared to 10 and 35 nM for paclitaxel in liposomes and Cremophor w 

EL-ethanol (50:50) formulations, respectively. 46 The incorporation of PEG-derivatized 



phospholipid is expected to enhance the in vivo circulation half life of the formulation, thereby 

enhancing the exposure of the drug to the targeted tumor mass. 

Another study reported the formulation of filter sterilizable emulsion formulation of paclitaxel 

using a-tocopherol as the oil phase and a-tocopherylpolyethyleneglycol-1,000 succinate (TGPS) 

and poloxamer 407 as emulsifiers. 47 The emulsion droplet size was O200 nm with high levels of 

paclitaxel loading (8–10 mg/mL). The formulation exhibited better efficacy and was more tolerable 

when studied in B16 melanoma tumor model in mice. In addition, another group of researchers 

developed a submicron emulsion (droplet diameter 150 nm) of paclitaxel with oil blend (triglycerols), 

egg PC, Tween w 80, and glycerol solution. The formulation showed cytotoxicity against 

HeLa cells with an IC 50 at 30 nM. An anti-tumor agent, valinomycin, was formulated in the 

emulsion form using the commercially-available Intralipid 10% soybean oil emulsion used in 

parenteral nutrition. Evaluation of this formulation in vivo indicated that the emulsion formulation 

produced similarly shaped dose-response curve to that of an aqueous suspension, but the emulsion 

formulation required a 30-fold lower dose than the suspension to produce therapeutic similar 

effects. 7 

In formulation of nanoemulsions, selection of the appropriate oil phase is important because 

most of the anti-cancer compounds exhibit poor solubility in the oil phase, especially those with 

highly lipophilic oils. Kan et al. determined the solubility of paclitaxel in various oils such as 

tributyrin, tricaproin, tricaprylin, corn, soyabean, cotton seed, and mineral oil, and they found that 

triacylglycerols with short fatty acid chains (tributyrin and tricaproin) were good solvents for 

paclitaxel with solubility of more than 9.00 mg/g as compared to other vegetable oils (range 

0.14–0.23 mg/gm). 48 Another approach to enhance the oil solubility of the anti-cancer compounds 

is the chemical modification or prodrug formation. Prodrugs with increased oil solubility have been 

obtained with such anti-cancer drugs as teniposide, etoposide, camptothecin, and paclitaxel, 

whereas amphiphilic derivatives have been prepared from fluorodeoxyuridine. 49 Esterification 

with long-chain fatty acids (i.e., oleic acid) has also been reported to increase the oil solubility 

of many anti-cancer drugs. 

Using submicron oil-in-water systems, one can engineer a multifunctional therapeutic system 

by combining several drugs in the same carrier (i.e., with an oil soluble drug in the core and an 

amphiphilic one in the surface monolayer) for more effective treatment of tumors or other diseased 

conditions. Emulsion formulations also show promise in cancer chemotherapy as vehicles for 

prolonging the drug release after intramuscular and intratumoral injection (W/O systems) and as 

a means of enhancing the transport of anti-cancer drugs via the lymphatic system. 10 

35.6.2 POSITIVELY CHARGED ANOEMULSIONS IN CANCER THERAPY 

Positively charged nanoemulsions systems are expected to interact with negatively charged cell 

surfaces more efficiently, and this aspect of the positively charged nanoemulsions has been 

explored for possibility of oligonucleotide delivery to cancer cells. 50–53 Because oligonucleotides 

molecules display a polyanioinc character and present a large molecular structure, their ability to 

cross cell membranes remains very low. Teixeira et al. 50 formulated cationic O/W emulsions of 

oligonucleotides (pdT 16) using medium chain triglycerides, PC, poloxamer, and either monocationic 

SA lipid PC–SA emulsion or a polycationic lipid, RPR C 18 (PC–RPRC 18 emulsion). They 

studied the biodistribution in leukemic P388/ADR cells grown in the peritoneal cavity after intraperitoneal 

administration in mice (Table 35.3). The fate of both the pdT and the marker of the oil 

phase (cholesteryl oleate) were determined in the fluid (devoid of cancer cells) and in the P388/ 

ADR cell pellet. The results indicated that pdT in solution remained only in the fluid and did not 

associate at all with the tumor cells. However, pdT injected as an emulsion formulation was 

detectable in the cancer cells pellet even after 24 h and in very high proportions (up to 18% of 

the injected dose). When the area-under-the-curve values of the concentration versus time profiles 

for different formulations were compared, it was observed that RPR C 18 formulations favored an 


732 

TABLE 35.3 

Area Under the Percent Tissues–Time Curve After Intraperitoneal Administration Into Mice 

Bearing P388/ADR Ascites of pdT 16 in Solution or Associated with PC/SA or PC/RPRC 18 

Emulsion 

AUC0–24 h (% Injected Dose/min) 

Preparations Blood Liver Kidneys Lungs Spleen Cell Pellet 

Solution 3,870 16,411 4,725 1,231 2,343 3,095 

PC/SA 3,393 16,505 4,214 1,254 2,438 13,634 

PC/RPRC 2,984 1,167 2,771 850 2,200 22,592 

PC/SA, phosphatidylcholine-stearylamine emulsion; PC/RPRC18, Posphatidylcholine-polycationic lipid RPRC18 emulsion; 

pdT16, Oligohexadecathymidylate. 

Source: Adapted from Teixeira, H., Dubernet, C., Chacun, H., Rabinovich, L., Boutet, V., Deverre, J. R., Benita, S., and 

Couvreur, P., J. Control Release, 89(3), 473–482, 2003. 

increased association of pdT to the tumor cells compared to SA (Table 35.3). Because the distribution 

of the oil marker in the tumor fluid and the P388/ADR cell pellet could not be correlated with 

that of pdT, the authors postulated that pdT was probably not taken through the endocytosis of the 

oil droplets but by positive charges of the emulsion that probably increased membrane permeability 

and allowed the pdT molecules to more efficiently enter the cells. In another study, cationic 

emulsions composed of 3b [N-(N 0 ,N 0 -DC-Chol)] and dioleoylphosphatidyl ethanolamine, castor 

oil, and Tween w 80 efficiently delivered plasmid DNA into various cancer cells with low toxicity. 

Furthermore, the use of chitosan as a condensing agent (for DNA) and subsequent complexation 

with cationic emulsion composed of DC-chol enhanced the transfection efficiency in vitro 

compared to DNA/emulsion complexes with the same formulation with chitosan. 54 In vivo study 

in mice showed that with chitosan enhanced emulsion complexes, the GFP mRNA expression was 

prolonged in liver and lung until day 6. Goldstein et al. 55 formulated a monoclonal antibody, 

AMB8LK-, cationic emulsion for targeted delivery to the tumors that overexpress H-ferritin 

(that is recognized by AMB8LK). The results of cell culture studies indicated that the coupling 

of AMB8LK-Fab’ fragment to the cationic emulsion increased the cells uptake by 50% as 

compared to non-Fab’ conjugated cationic emulsion. Similar conjugation of B-cell lymphoma 

monoclonal antibody, LL2, to the surface of polyethylene glycol-PE-based lipid emulsion globules 

has been reported with improved binding to the B-lymphoma cells. 56 The results from these studies 

indicate significant potential of cationic emulsion systems as effective drug/gene delivery vehicles 

to tumor cells. 

35.6.3 PHOTODYNAMIC THERAPY OF CANCER 


Photodynamic therapy (PDT) of cancer is based on the concept that certain photosensitizers can be 

localized in the neoplastic tissue, and subsequently, these photosensitizers can be activated with the 

appropriate wavelength (energy) of light to generate active molecular species such as free radicals 

and singlet oxygen ( 1 O2) that are toxic to cells and tissues. 57–59 PDT is a binary therapy, and it has a 

potential advantage in its inherent dual selectivity. First, selectivity is achieved by an increased 

concentration of the photosensitizer in target tissue; second, the irradiation can be limited to a 

specified volume. Provided that the photosensitizer is nontoxic, only the irradiated areas will be 

affected even if the photosensitizer does bind to normal tissues. Selectivity can be further enhanced 

by binding photosensitizers to molecular delivery systems that have high affinity for target tissue. 

Many photosensitizers, however, are characterized by a high level of hydrophobicity. The poor 

water solubility of most photosensitizers often prevents their direct administration into the 



bloodstream. To overcome this problem, photosensitizers have to be either chemically modified or 

incorporaed into lipid delivery systems such as liposomes and emulsions. 

Morgan et al. utilized emulsions prepared with the ethoxy castor oil Cremophor w -EL (CR) and 

DPPC liposomes for the administration of several hydrophobic purpurins to rats with an implanted 

urothelial tumor. 60 PDT studies showed that tin (IV) etiopurpurin (SnET2) at a dose of 1 mg 

kg-winduced 100% tumor cure with both delivery vehicles. At lower doses, SnET2 in CR was 

more effective than SnET2 in liposomes. 61 Therapeutic doses of SnET2 in CR caused no phototoxicity 

to footpad tissue irradiated 24 h after administration although some phototoxicity was 

observed with liposomes. A higher PDT efficacy was observed by delivering a ketochlorin in 

CR than in Tween w 80. The better response of the RIF tumor (murine mammary sarcoma) to 

PDT was ascribed to an increased tumor uptake of the ketochlorin, probably caused by a longer 

persistence of the drug in the circulation. 62 Similar conclusions were drawn by comparing the 

pharmacokinetics of several germanium (IV)-octabutoxy–phthalocyanines (GePc) incorporated 

into DPPC liposomes or in CR emulsion. The Cremophor-delivered GePcs were cleared from 

the blood circulation at a much slower rate than the liposome-delivered GePcs. 63 At the same 

time, Cremophor induced a slower and reduced uptake of the GePcs in the liver and spleen, and it 

greatly enhanced the uptake in the tumor as compared to liposomes. Furthermore, CR induced a 

four-fold higher accumulation of the Pcs in the MS-2 fibrosarcoma, yielding in all cases two to three 

times higher tumor-to-muscle (peritumoral tissue) and tumor-to-skin drug concentration ratios. 64 

Almost similar results have been reported with silicon (IV)-phthalocyanine (SiPc) CR or liposome 

formulation in C57B1/6 mice Lewis lung carcinoma. 65 

In brief, various PDT therapies have reported two different vehicles for photosensitizers, a 

Cremophor oil emulsion and DPPC liposomal vesicles. The reported pharmacokinetic studies 

clearly indicate that the former vehicle yields a significantly larger selectivity of tumor targeting, 

mainly as a consequence of an enhanced accumulation in the malignant lesion, and no appreciable 

differences in the uptake by healthy tissues is observed with Cremophor or DPPC liposomes (except 

higher uptake by brain tissues with CR formulation). The greater accumulation of Cremophordelivered 

photosensitizers has been thought to be attributed to the tendency of this vehicle to release 

the associated photosensitizing drug to serum low-density lipoproteins (LDLs) in a highly preferential 

amount, and DPPC liposomes transfer the drug to all the components of the lipoprotein class 

in aliquots that are proportional to the relative concentration of the individual proteins. 61 LDLs are 

known to display a preferential interaction with a variety of neoplastic cells through a receptormediated 

endocytotic process. The tumor cells often express an increased number of LDL receptors 

as compared with normal cells. 

35.6.4 NEUTRON CAPTURE THERAPY OF CANCER 

Neutron Capture Therapy (NCT) is a binary radiation therapy modality that brings together two 

components that when kept separate, have only minor effects on the cells. The first component is a 

stable isotope of boron or gadolinium (Gd) that can be concentrated in tumor cells by a suitable 

delivery vehicle. The second is a beam of low-energy neutrons. Boron or Gd in or adjacent to the 

tumor cells disintegrates after capturing a neutron, and the high energy heavy charged particles 

produced through this interaction destroy only the cancer cells in close proximity to it, leaving 

adjacent normal cells largely unaffected. 66 The success of NCT relies on the targeting of boron and 

Gd-based compounds to the tumor mass and to achieve desirable intracellular concentrations of 

these agents. At the present time, there are two targets with NCT, namely glioblastoma (malignant 

brain tumor) and malignant melanoma. 

Lu and co-workers developed and evaluated a very low-density lipoprotein (VLDL), resembling 

phospholipid-submicron emulsion as a carrier system for new cholesterol-based boronated 

compound, BCH (anti-cancer boron neutron capture therapy compound), for targeted delivery to 

cancer cells. 67 Emulsions were formulated by sonication method using triolein, cholesterol, 


734 

cholesterol oleate as lipids, egg PC, lysophosphatidylcholine as surfactants, and BCH as active 

pharmaceutical ingredient. The BCH was solubilized in the inner oily core of the emulsion droplets 

(155G5 nm). The formulation was designed based on the fact that LDL and high-density lipoproteins 

(HDL) are natural carriers of cholesteryl esters in the body, and certain human and animal 

tumor types have been shown to have elevated LDL-receptor activity primarily because the rapidly 

dividing cancer cells require higher amounts of cholesterol to build new cell membranes. It is 

expected that VLDL-resembling formulation may mimic the VLDL–LDL biological process for 

targeted drug delivery to cancer cells. Cell culture data showed sufficient uptake of BCH in rat 9L 

glioma cells (O50 mg boron/g cells). The authors proposed that such VLDL-resembling formulation 

may serve as a targeted drug delivery system to cancerous cells in vivo. 

Miyamoto et al. prepared the nanoemulsions of Gd complexed with DTPA-distearylamide (SA) 

and DTPA-distearylester (SE) using soybean oil and hydrogenated PC and various co-surfactants 

(Cremophore w -RH60, Myrj w 53, Myrj w 59, or Brij w 700). 68,69 The biodistribution of the Gd was 

studied in melanoma-bearing hamsters after intraperitoneal administration of the formulations. The 

co-surfactant-containing emulsions were more rapidly and completely cleared from the abdominal 

cavity than the plain emulsion without co-surfactants. When the effect of the co-surfactants on the 

biodistribution of Gd from Gd–DTPA–SA-containing emulsions in the standard-Gd formulation 

were compared, the Cremophore w -RH60 emulsion exhibited the highest Gd accumulation in the 

tumor (107 mg Gd/g tumor), possibly resulting from its small particle size (78 nm) and the stable 

coat on the particle surfaces with PEO. The Brij w 700-containing emulsion allowed the highest 

blood Gd concentration for a prolonged period. However, it exhibited slower Gd accumulation in 

the tumor, only reaching an identical level in comparison with the Cremophore w -RH60 emulsion. 

Gd–DTPA–SE emulsions exhibited poor tumor accumulation of Gd. 

The same group of researchers evaluated the effect of intravenous administration of Gd-nanoemulsions 

on biodistribution and tumor accumulation. 70 The biodistribution data revealed that 

the intravenous injection had three advantages over the ip injection, namely, a faster and higher 

accumulation of Gd and a more extended retention time of Gd. Two intravenous injections of the 

standard-Gd-nanoemulsion (1.5 mg Gd/ml) with a 24 h interval doubled the tumor accumulation of 

Gd, resulting in 49.7 mg Gd/g wet tumor 12 h after administration. By using a two-fold Gd-enriched 

formulation (3.0 mg Gd/ml) in the repeated administration schedule, the accumulation was 

doubled again, reaching 101 mg Gd/g wet tumor. This level was comparable to the maximum level 

in the single ip injection previously reported. The authors proposed that intravenous injection could 

be an alternative to ip injection as an administration route for Gd. 

35.6.5 PERFLUOROCHEMICAL NANOEMULSIONS IN CANCER THERAPY 


Perfluorocarbon emulsions are being clinically evaluated as artificial oxygen carriers to reduce 

allogeneic blood transfusions or to improve tissue oxygenation. 13 Perfluorochemicals are chemically 

inert synthetic molecules that primarily consist of carbon and fluorine atoms and are clear, colorless 

liquids. They have the ability to physically dissolve significant quantities of many gases, including 

oxygen and carbon dioxide. Perfluorochemicals are hydrophobic and are not miscible with water. 

Perfluorochemicals have to be emulsified for intravenous use. To mimic the natural oxygen carrying 

cells (RBCs), the droplet size of perfluorocarbon emulsions is maintained in sub-micron range 

(median diameter !0.2 mm). Egg phospholipid has been used as an emulsifier of choice in these 

formulations. The examples of the commercial perfluorocarbon emulsions are Oxygente (Alliance 

Pharmaceutical Corporation, San Diego, CA, U.S.A.), Oxyfluor w (Hemagen Inc., St Louis, MO, 

U.S.A.) and Fluosol-DA (Alpha Therapeutic Corp., Los Angeles, CA, U.S.A.). 

The perfluorochemical nanoemulsions (PFCE) have opened interesting opportunities in cancer 

therapy. It is suggested that fluorocarbon emulsions might find a role in photodynamic therapy, both 

as carriers for sensitizing dyes and also to maintain tissue oxygenation in hypoxic regions of solid 

tumors. The high solubility of oxygen in fluorocarbon emulsions maintains solution oxygen 



tension, optimizing photo-oxidative damage. The hydrophobic anti-cancer drugs can be delivered 

to the tumor mass by dissolving them in a hydrophobic core of the emulsion. Furthermore, PFCE 

can be used as an adjuvant to radiation therapy and/or chemotherapy in the treatment of solid 

tumors. 71,72 The rationale for the use of PFCE in this therapeutic setting is that solid tumor masses 

contain areas of hypoxia that are therapeutically resistant. Since x-rays and many chemotherapeutic 

agents require oxygen to be maximally cytotoxic and most normal tissues are well-oxygenated, the 

additional oxygen put in circulation by the PFCE should not increase the normal tissue toxicities 

produced by the various therapies. The preclinical studies have shown very positive effects with 

single dose and fractionated radiation in several rodent solid tumor models. Many widely used anticancer 

drugs, including anti-tumor alkylating agents and doxorubicin, have shown improved 

response by PFCE coadministration. 73 Also, local application of toxic doses of PFCEs resulted 

in the necrosis of cancer cells. This is especially promising in the treatment of cancers of the head 

and neck regions that are currently difficult to treat. 74 

Lanza and co-workers developed a perfluorocarbon emulsion that can be used for ultrasound 

and magnetic resonance imaging (MRI) detection of fibrin and thrombus. 75 Liquid perfluorocarbon 

inside the nanoemulsion core makes it stable and acoustically reflective—a benefit for ultrasound 

imaging. For MRI, the nanoemulsion can carry Gd molecules. The authors estimates that one 

nanoemulsion droplet can carry up to 100,000 Gd molecules, enough for a large MR signal and 

effective imaging of the targeted area. The fluorine component allows quantification of drug 

amounts to the targeted area using MR spectroscopy. The authors speculated that one could 

attach radionuclides such as technetium-99m for nuclear imaging or radio-opaque compounds 

for contrast enhancement in computed tomography in the nanoemulsion core. The same group 

of researchers have also developed an avb3 integrin-targeted paramagnetic nanoemulsions that can 

detect and characterize sparse integrin expression on tumor neovasculature induced by nascent 

melanoma xenografts. 76 Alphan-beta3 integrin is an excellent marker for angiogenesis and, therefore, 

is used as the target to dock the nanoemulsion particles. When the emulsion particles are 

injected into the body, the homing ligand causes the droplets to seek out a vb 3 integrin. Detecting 

a vb 3 integrin provides a direct and early measure of tumor vascularity. Angiogenesis is associated 

with a wide variety of pathologies, and by targeting angiogenesis, the nanoemulsions can identify 

solid tumors and lymphoid cancers at an early stage. Because angiogenesis also occurs with 

atherosclerosis and inflammatory diseases, authors feel that such nanoemulsions can be used for 

additional diagnostic applications, including imaging early signs of atherosclerosis, restenosis 

following angioplasty, and inflammatory diseases such as rheumatoid arthritis. 


Nanoemulsion formulations offer several advantages for the delivery of drugs, biologicals, or 

diagnostic agents. Traditionally, nanoemulsions have been used in clinics for more than four 

decades as total parenteral nutrition fluids. Several other products for drug delivery applications 

such as Diprivan w , Liple w , and Ropion w have also reached the marketplace. Although nanoemulsions 

are chiefly seen as vehicles for administering aqueous insoluble drugs, they have more 

recently received increasing attention as colloidal carriers for targeted delivery of various anticancer 

drugs, photosensitizers, neutron capture therapy agents, or diagnostic agents. Because of 

their sub-micron size, they can be easily targeted to the tumor area. Moreover, the possibility of 

surface functionalization with a targeting moiety has opened new avenues for targeted delivery of 

drugs, genes, photosensitizers, and other molecules to the tumor area. Research with perflurochemical 

nanoemulsions has shown promising results for the treatment of cancer in conjugation with 

other treatment modalities and targeted delivery to the neovasculature. It is expected that further 

research and development work will be carried out in the near future for clinical realization of these 

targeted delivery vehicles. 


736 

REFERENCES 


1. Salager, J. L., Formulation concepts for the emulsion maker, In Pharmaceutical Emulsions and 

Suspensions, Nielloud, F. and Marti-Mestres, G., Eds., Marcel Dekker, New York, pp. 19–72, 2000. 

2. Tenjarla, S., Microemulsions: An overview and pharmaceutical applications, Crit. Rev. Ther. Drug 

Carrier Syst., 16(5), 461–521, 1999. 

3. Keipert, S., Siebenbrodt, I., Luders, F., and Bornschein, M., Microemulsions and their potential 

pharmaceutical application, Pharmazie, 44(7), 433–444, 1989. 

4. Schoot, H., Colloidal dispersions, In Remington: The Science and Practice of Pharmacy, Gennaro, A., 

Lippincott, R., Ed., 20th ed., Williams & Witkins, Maryland, pp. 288–315, 2000. 

5. Sarker, D. K., Engineering of nanoemulsions for drug delivery, Curr. Drug Deliv., 2(4), 297–310, 

2005. 

6. Jores, K., Lipid nanodispersions as drug carrier systems -a physicochemical characterization, http:// 

sundoc.bibliothek.uni-halle.de/diss-online/04/04H310/prom.pdf, accessed date, Aug. 16, 2005. 

7. Buszello, K. and Muller, B., Emulsions as drug delivery systems, In Pharmaceutical Emulsions and 

Suspensions, Nielloud, F. and Marti-Mestres, G., Eds., Marcel Dekker, New York, pp. 191–228, 2000. 

8. Yang, S. C. and Benita, S., Enhanced absorption and drug targeting by positively charged submicron 

emulsions, Drug Dev. Res., 50, 476–486, 2000. 

9. Inactive ingredient search for approved drug products, http://www.accessdata.fda.gov/scripts/cder/ 

iig/index.cfm, date accessed, Nov. 05, 2005. 

10. Eccleston, G., Emulsions, In Encyclopedia of Pharmaceutical Technology, Swarbrick, J. and Boylan, 

J., Eds., Marcel Dekker, New York, pp. 137–188, 1992. 

11. Chung, H., Kim, T. W., Kwon, M., Kwon, I. C., and Jeong, S. C., Oil components modulate physical 

characteristics and function of the natural oil emulsions as drug or gene delivery system, J. Control 

Release, 71, 339–350, 2001. 

12. Allison, A. D., Squelene and squalane emulsions as adjuvants, Methods, 19, 87–93, 1999. 

13. Lane, T. A., Perfluorochemical-based artificial oxygen carrying red cell substitutes, Transfus. Sci., 

16(1), 19–31, 1995. 

14. Benita, S. and Levy, M. Y., Submicron emulsions as colloidal drug carrieres for intravenous administration: 

Comprehensive physicochemical characterization, J. Pharm. Sci., 82, 1069–1079, 1993. 

15. Klang, S. and Benita, S., Design and evaluation of submicron emulsions as colloidal drug carriers for 

intravenous administration, In Submicron Emulsions in Drug Targeting and Delivery, Benita, S., Ed., 

Academic Publishers, Harwood, pp. 119–152, 1998. 

16. Klang, S., Frucht-Pery, J., Hoffman, A., and Benita, S., Physicochemical characterization and acute 

toxicity evaluation of a positively charged submicron emulsion vehicle, J. Pharm. Pharmacol., 46, 

986–993, 1998. 

17. Haskell, R., Nanotechnology for drug delivery, www.banyu-zaidan.or.jp/symp/soyaku/haskell.pdf, 

date accessed, Oct. 06, 2005. 

18. Haskell, R. J., Characterization of submicron systems via optical methods, J. Pharm. Sci., 87(2), 

125–129, 1998. 

19. Wabel, C., Influence of lecithin on structure and stability of parenteral fat emulsions, http://www2. 

chemie.uni-erlangen.de/services/dissonline/data/dissertation/Christoph_Wabel/html/index.html, 

accessed date Oct. 09, 2005. 

20. Salager, J. L., Emulsion properties and relate know-how and attain them, In Pharmaceutical Emulsions 

and Suspensions, Nielloud, F. and Marti-Mestres, G., Eds., Marcel Dekker, New York, pp. 73–125, 2000. 

21. Gaysorn, C., Lyons, R. T., Patel, M. V., and Hem, S. L., Effect of surface charge on the stability of 

oil/water emulsions during steam sterilization, J. Pharm. Sci., 88(4), 454–458, 1999. 

22. Elbaz, E., Zeevi, A., Klang, S., and Benita, S., Positively charged submicron emulsions- a new type of 

colloidal drug carrier, Int. J. Pharm., 96, R1–R6, 1993. 

23. Klang, S. H., Siganos, C. S., Benita, S., and Frucht-Pery, J., Evaluation of a positively charged 

submicron emulsion of piroxicam on the rabbit corneum healing process following alkali burn, 

J. Control Release, 57(1), 19–27, 1999. 

24. Klang, S. H., Frucht-Pery, J., Hoffman, A., and Benita, S., Physicochemical characterization and acute 

toxicity evaluation of a positively-charged submicron emulsion vehicle, J. Pharm. Pharmacol., 

46(12), 986–993, 1994. 



25. Gershanik, T., Benzeno, S., and Benita, S., Interaction of a self-emulsifying lipid drug delivery system 

with the everted rat intestinal mucosa as a function of droplet size and surface charge, Pharm. Res., 

15(6), 863–869, 1998. 

26. Oku, N., Tokudome, Y., Namba, Y., Saito, N., Endo, M., Hasegawa, Y., Kawai, M., Tsukada, H., and 

Okada, S., Effect of serum protein binding on real-time trafficking of liposomes with different charges 

analyzed by positron emission tomography, Biochim. Biophys. Acta, 1280(1), 149–154, 1996. 

27. Gao, X. and Huang, L., A novel cationic liposome reagent for efficient transfection of mammalian 

cells, Biochem. Biophys. Res. Commun., 179(1), 280–285, 1991. 

28. Calvo, P., Alonso, M. J., Vila-Jato, J. L., and Robinson, J. R., Development of positively charged 

colloidal drug carrieres: Chitosan coated polyester nanocapsules and sub-micron emulsions, Colloid 

Polm. Sci., 275, 46–53, 1997. 

29. Jumaa, M. and Muller, B. W., Physicochemical properties of chitosan-lipid emulsions and their 

stability during the autoclaving process, Int. J. Pharm., 183(2), 175–184, 1999. 

30. Samama, J. P., Lee, K. M., and Biellmann, J. F., Enzymes and microemulsions: Activity and kinetic 

properties of liver alcohol dehydrogenase in ionic water in oil microemulsions, Eur. J. Biochem., 163, 

609–617, 1987. 

31. Kurihara, A., Shibayama, Y., Mizota, A., Yasuno, A., Ikeda, M., and Hisaoka, M., Pharmacokinetics 

of highly lipophilic antitumor agent palmitoyl rhizoxin incorporated in lipid emulsions in rats, Biol. 

Pharm. Bull., 19(2), 252–258, 1996. 

32. Lundberg, B. B., Preparation of drug carrier emulsions stabilized with phosphatidylcholine surfactant 

mixtures, J. Pharm. Sci., 83, 72–75, 1994. 

33. Liu, F. and Liu, D., Long-circulating emulsions (oil-in-water) as carriers for lipophilic drugs, Pharm. 

Res., 12(7), 1060–1064, 1995. 

34. Hodoshima, N., Udagawa, C., Ando, T., Fukuyasu, H., Watanabe, T., and Nakabayashi, S., Lipid 

nanoparticles for delivering antitumor drugs, Int. J. Pharm., 146, 81–92, 1997. 

35. Kurihara, A., Shibayama, Y., Mizota, A., Yasuno, A., Ikeda, M., Sasagawa, K., Kobayashi, T., and 

Hisaoka, M., Enhanced tumor delivery and antitumor activity of palmitoyl rhizoxin using stable lipid 

emulsions in mice, Pharm. Res., 13(2), 305–310, 1996. 

36. Davis, S. S., Washington, C., West, P., Illum, L., Liversidge, G., Sternson, L., and Kirsh, R., Lipid 

emulsions as drug delivery systems, Ann. N.Y. Acad. Sci., 507, 75–88, 1987. 

37. Arimoto, I., Matsumoto, C., Tanaka, M., Okuhira, K., Saito, H., and Handa, T., Surface composition 

regulates clearance from plasma and triolein lipolysis of lipid emulsions, Lipids, 33(8), 773–779, 

1998. 

38. Krobbuaban, B., Diregpoke, S., Kumkeaw, S., and Tanomsat, M., Comparison on pain on injection of 

a small particle size-lipid emulsion of propofol and standard propofol with or without lidocaine, 

J. Med. Assoc. Thai., 88(10), 1401–1405, 2005. 

39. Thorn-Alquist, A. M., Parenteral use of diazepam in an emulsion formulation. A clinical study, Acta 

Anaesthesiol. Scand., 21(5), 400–404, 1977. 

40. Westrin, P., Jonmarker, C., and Werner, O., Dissolving methohexital in a lipid emulsion reduces pain 

associated with intravenous injection, Anesthesiology, 76(6), 930–934, 1992. 

41. Lovell, M. W., Johnson, H. W., Hui, H. W., Cannon, J. B., Gupta, P. K., and Hsu, C. C., Less-painful 

emulsion formulations for intravenous administration of clarithomycin, Int. J. Pharm., 109, 45–57, 

1994. 

42. Stuttmann, H., Doenicke, A., Kugler, J., and Laub, M., A new formulation of etomidate in lipid 

emulsion: Bioavailability and venous irritation, Anaesthesist, 38, 421–423, 1989. 

43. Glen, J. B. and Hunter, S. C., Pharmacology of an emulsion formulation of ICI 35 868, Br. 

J. Anaesth., 56(6), 617–626, 1984. 

44. Jang, S. H., Wientjes, M. G., Lu, D., and Au, J. L., Drug delivery and transport to solid tumors, Pharm. 

Res., 20(9), 1337–1350, 2003. 

45. Mayhew, E. G., Lasic, D., Babbar, S., and Martin, F. J., Pharmacokinetics and antitumor activity of 

epirubicin encapsulated in long-circulating liposomes incorporating a polyethylene glycol-derivatized 

phospholipid, Int. J. Cancer, 51(2), 302–309, 1992. 

46. Lundberg, B. B., A submicron lipid emulsion coated with amphipathic polyethylene glycol for 

parenteral administration of paclitaxel (Taxol), J. Pharm. Pharmacol., 49(1), 16–21, 1997. 


738 


47. Constantinides, P. P., Lambert, K. J., Tustian, A. K., Schneider, B., Lalji, S., Ma, W., Wentzel, B., 

Kessler, D., Worah, D., and Quay, S. C., Formulation development and antitumor activity of a 

filter-sterilizable emulsion of paclitaxel, Pharm. Res., 17(2), 175–182, 2000. 

48. Kan, P., Chen, Z. B., Lee, C. J., and Chu, I. M., Development of nonionic surfactant/phospholipid o/w 

emulsion as a paclitaxel delivery system, J. Control Release, 58(3), 271–278, 1999. 

49. Lundberg, B. B., Sub-micron lipid emulsions as carrieres for anticancer drugs, http://www.cmbl.org. 

pl/lundberg.pdf, accessed date, Sept. 25, 2005. 

50. Teixeira, H., Dubernet, C., Chacun, H., Rabinovich, L., Boutet, V., Deverre, J. R., Benita, S., and 

Couvreur, P., Cationic emulsions improves the delivery of oligonucleotides to leukemic P388/ADR 

cells in ascite, J. Control Release, 89(3), 473–482, 2003. 

51. Teixeira, H., Dubernet, C., Puisieux, F., Benita, S., and Couvreur, P., Submicron cationic emulsions as 

a new delivery system for oligonucleotides, Pharm. Res., 16(1), 30–36, 1999. 

52. Teixeira, H., Dubernet, C., Rosilio, V., Laigle, A., Deverre, J. R., Scherman, D., Benita, S., and 

Couvreur, P., Factors influencing the oligonucleotides release from O-W submicron cationic emulsions, 

J. Control Release, 70(1–2), 243–255, 2001. 

53. Teixeira, H., Rosilio, V., Laigle, A., Lepault, J., Erk, I., Scherman, D., Benita, S., Couvreur, P., and 

Dubernet, C., Characterization of oligonucleotide/lipid interactions in submicron cationic emulsions: 

Influence of the cationic lipid structure and the presence of PEG-lipids, Biophys. Chem., 92(3), 

169–181, 2001. 

54. Lee, M. K., Chun, S. K., Choi, W. J., Kim, J. K., Choi, S. H., Kim, A., Oungbho, K., Park, J. S., Ahn, 

W. S., and Kim, C. K., The use of chitosan as a condensing agent to enhance emulsion-mediated gene 

transfer, Biomaterials, 26(14), 2147–2156, 2005. 

55. Goldstein, D., Nassar, T., Lambert, G., Kadouche, J., and Benita, S., The design and evaluation of a 

novel targeted drug delivery system using cationic emulsion-antibody conjugates, J. Control Release, 

108(2–3), 418–432, 2005. 

56. Lundberg, B. B., Griffiths, G., and Hansen, H., Conjugation of an anti-B-cell lymphoma monoclonal 

antibody, LL2, to long-circulating drug-carrier lipid emulsions, J. Pharm. Pharmacol., 51(10), 

1099–1105, 1999. 

57. Waldow, S. M., Henderson, B. W., and Dougherty, T. J., Hyperthermic potentiation of photodynamic 

therapy employing Photofrin I and II: Comparison of results using three animal tumor models, Lasers 

Surg. Med., 7(1), 12–22, 1987. 

58. Dougherty, T. J., Photosensitizers: therapy and detection of malignant tumors, Photochem. Photobiol., 

45(6), 879–889, 1987. 

59. Potter, W. R., Mang, T. S., and Dougherty, T. J., The theory of photodynamic therapy dosimetry: 

Consequences of photo-destruction of sensitizer, Photochem. Photobiol., 46(1), 97–101, 1987. 

60. Morgan, A. R., Garbo, G. M., Keck, R. W., and Selman, S. H., New photosensitizers for photodynamic 

therapy: Combined effect of metallopurpurin derivatives and light on transplantable bladder 

tumors, Cancer Res., 48(1), 194–198, 1988. 

61. Reddi, E., Role of delivery vehicles for photosensitizers in the photodynamic therapy of tumours, 

J. Photochem. Photobiol. B, 37(3), 189–195, 1997. 

62. Woodburn, K., Chang, C. K., Lee, S., Henderson, B., and Kessel, D., Biodistribution and PDT efficacy 

of a ketochlorin photosensitizer as a function of the delivery vehicle, Photochem. Photobiol., 60(2), 

154–159, 1994. 

63. Soncin, M., Polo, L., Reddi, E., Jori, G., Kenney, M. E., Cheng, G., and Rodgers, M. A., Effect of the 

delivery system on the biodistribution of Ge(IV) octabutoxy-phthalocyanines in tumour-bearing mice, 

Cancer Lett., 89(1), 101–106, 1995. 

64. Soncin, M., Polo, L., Reddi, E., Jori, G., Kenney, M. E., Cheng, G., and Rodgers, M. A., Effect of axial 

ligation and delivery system on the tumour-localising and -photosensitising properties of Ge(IV)octabutoxy-phthalocyanines, 

Br. J. Cancer, 71(4), 727–732, 1995. 

65. Wohrle, D., Muller, S., Shopova, M., Mantareva, V., Spassova, G., Vietri, F., Ricchelli, F., and Jori, 

G., Effect of delivery system on the pharmacokinetic and phototherapeutic properties of bis(methyloxyethyleneoxy) 

silicon-phthalocyanine in tumor-bearing mice, J. Photochem. Photobiol. B, 

50(2–3), 124–128, 1999. 

66. The basics of boron neutron capture therapy, http://web.mit.edu/nrl/www/bnct/info/description/ 

description.html, date accessed, Nov. 27, 2005. 



67. Shawer, M., Greenspan, P., OIe, S., and Lu, D. R., VLDL-resembling phospholipid-submicron emulsion 

for cholesterol-based drug targeting, J. Pharm. Sci., 91(6), 1405–1413, 2002. 

68. Miyamoto, M., Hirano, K., Ichikawa, H., Fukumori, Y., Akine, Y., and Tokuuye, K., Biodistribution 

of gadolinium incorporated in lipid emulsions intraperitoneally administered for neutron-capture 

therapy with tumor-bearing hamsters, Biol. Pharm. Bull., 22(12), 1331–1340, 1999. 

69. Miyamoto, M., Hirano, K., Ichikawa, H., Fukumori, Y., Akine, Y., and Tokuuye, K., Preparation of 

Gadolinium-containing emulsions stabilized with phosphatidylcholine surfactant mixtures for neutron 

capture therapy, Chem. Pharm. Bull. (Tokyo), 47(2), 203–208, 1999. 

70. Watanabe, T., Ichikawa, H., and Fukumori, Y., Tumor accumulation of gadolinium in lipid-nanoparticles 

intravenously injected for neutron-capture therapy of cancer, Eur. J. Pharm. Biopharm., 54(2), 

119–124, 2002. 

71. Teicher, B. A., Use of perfluorochemical emulsions in cancer therapy, Biomater. Artif. Cells Immobil. 

Biotechnol., 20(2–4), 875–882, 1992. 

72. Teicher, B. A., Herman, T. S., and Frei, E., Perfluorochemical emulsions: Oxygen breathing in 

radiation sensitization and chemotherapy modulation, Important Adv. Oncol., 39–59, 1992. 

73. Teicher, B. A., Holden, S. A., Ara, G., Ha, C. S., Herman, T. S., and Northey, D., A new concentrated 

perfluorochemical emulsion and carbogen breathing as an adjuvant to treatment with antitumor 

alkylating agents, J. Cancer Res. Clin. Oncol., 118(7), 509–514, 1992. 

74. Rockwell, S., Irvin, C. G., and Nierenburg, M., Effect of a perfluorochemical emulsion on the 

development of artificial lung metastases in mice, Clin. Exp. Metastasis, 4(1), 45–50, 1986. 

75. Sisk, J., Imaging cancer one nanometer at a time, http://www.radiologytoday.net/imagingcancer.htm, 

accessed date, Sept. 06, 2005. 

76. Schmieder, A. H., Winter, P. M., Caruthers, S. D., Harris, T. D., Williams, T. A., Allen, J. S., and 

Lacy, E. K., Molecular MR imaging of melanoma angiogenesis with alphanubeta-3-targeted paramagnetic 

nanoparticles, Magn. Reson. Med., 53(3), 621–627, 2005. 


36 

CONTENTS 

Solid Lipid Nanoparticles for 

Anti-Tumor Drug Delivery 

Ho Lun Wong, Yongqiang Li, Reina Bendayan, 

Mike Andrew Rauth, and Xiao Yu Wu 

36.1 Introduction ....................................................................................................................... 742 

36.2 Composition and Structure of SLN .................................................................................. 743 

36.2.1 Lipids Used in SLN ............................................................................................ 744 

36.2.2 Surfactants Used in SLN..................................................................................... 747 

36.2.3 Other Agents Used in SLN ................................................................................. 748 

36.2.4 Types, Proposed Structure, and Morphology of SLN ........................................ 748 

36.2.4.1 Classical SLN ..................................................................................... 748 

36.2.4.2 Nanostructured Lipid Carrier (NLC) ................................................. 749 

36.2.4.3 Polymer–Lipid Hybrid Nanoparticles (PLN) .................................... 749 

36.3 Manufacturing Methods .................................................................................................... 751 

36.3.1 High Pressure Homogenization .......................................................................... 751 

36.3.2 Microemulsion..................................................................................................... 751 

36.3.3 Solvent Evaporation ............................................................................................ 753 

36.3.4 w/o/w Double Emulsion Method ........................................................................ 753 

36.4 Characteristic Properties of SLN ...................................................................................... 753 

36.4.1 Drug Loading ...................................................................................................... 753 

36.4.2 Drug Release ....................................................................................................... 754 

36.4.3 Stability of SLN .................................................................................................. 755 

36.5 SLN for Anti-Tumor Drug Delivery................................................................................. 756 

36.5.1 Cancer Chemotherapy ......................................................................................... 756 

36.5.2 Delivery of Cytotoxic Compounds Using SLN.................................................. 757 

36.5.3 Poorly Water-Soluble Compounds ..................................................................... 759 

36.5.3.1 General Principle................................................................................ 759 

36.5.3.2 Topoisomerase I Inhibitors (Camptothecin, Irinothecan) ................. 759 

36.5.3.3 Paclitaxel ............................................................................................ 760 

36.5.3.4 Other Poorly Water-Soluble Anti-Tumor Compounds 

(Etoposide, All-Trans Retinoic Acid, Butyric Acid)......................... 761 

36.5.4 Water-Soluble Ionic Salts ................................................................................... 761 

36.5.4.1 General Principle................................................................................ 761 

36.5.4.2 Doxorubicin and Idarubicin ............................................................... 761 

36.5.4.3 Water-Soluble, Non-Ionic Drug Molecule ........................................ 762 

36.5.4.3.1 General Principle ........................................................... 762 

36.5.4.3.2 Fluorouracil .................................................................... 763 

36.5.4.3.3 Overall Remarks on Cytotoxic Drug Delivery ............. 763 


741

742 

36.5.5 Delivery of Chemosensitizing Agents Using SLN............................................. 764 

36.5.5.1 Multidrug Resistance and Chemosensitizers ..................................... 764 

36.5.5.2 Rationales for Delivery of Chemosensitizers Using SLN................. 764 

36.5.5.3 Current Use of SLN for Chemosensitizer Delivery .......................... 765 

36.5.5.3.1 Verapamil and Quinidine............................................... 765 

36.5.5.3.2 Elacridar (GG918) ......................................................... 765 

36.5.5.3.3 Overall Remarks on Chemosensitizer Delivery ............ 766 

36.5.6 Effect of SLN Formulation on In Vitro Anti-Tumor Toxicity .......................... 766 

36.5.6.1 In Vivo Efficacy of SLN Delivered Antineoplastic Agents .............. 768 

36.5.6.2 Improved Pharmacokinetics and Drug Distribution .......................... 768 


References .................................................................................................................................... 771 



Solid lipid nanoparticles (SLN) are colloidal particles of a lipid matrix that is solid at body 

temperature. Since their first introduction by Müller et al. in 1991, 1 SLN have attracted increasing 

interest as a carrier system for therapeutic and cosmetic applications. 2–7 As a drug delivery system, 

SLN have been investigated in the last ten years for pharmacological and dermatological formulation 

development. They can be administered through a number of routes including parenteral, 

peroral, dermal and rectal. 2,4,6,7 Improved bioavailability and targeting capacity have been 

observed 8,9 and enhanced cytotoxicity against multidrug resistant cancer cells 10,11 have been 

evidenced when SLN are used as the delivery vehicles. 

SLN have been proposed as an alternative to other controlled drug delivery systems (CDDS) such 

as lipid emulsion, liposome, and polymeric nanoparticles as a result of their several advantages. For 

instance, in comparison to lipid emulsion, the solid lipid matrix of SLN makes sustained drug release 

possible. The solid lipids also immobilize drug molecules, thereby protecting the labile and sensitive 

drugs from coalescence and degradation, 12 and reduce drug leakage that are commonly seen in many 

other CDDS such as liposomes. Compared with some polymeric nanoparticles, SLN are generally 

less toxic because physiological and biocompatible lipids are utilized. Meanwhile, all of the less toxic 

surfactants that have been applied to other CDDS are equally applicable for SLN preparation. Other 

appealing features of SLN include the feasibility for mass production, 13–16 flexibility in sterilization, 4 

and avoidance of organic solvents in a typical SLN preparation process. It should be noted that SLN 

are also a versatile formulation. Both lipophilic and hydrophilic compounds can be encapsulated and 

delivered by SLN with modification in the formulation. 

The aforementioned useful qualities of SLN make them particularly attractive for the delivery 

of cancer chemotherapeutic agents. Anti-tumor drugs, especially the cytotoxic compounds that are 

used in conventional chemotherapy, are unique when compared to other classes of drugs in a 

number of areas such as the strong toxicity that is typical of cytotoxic drugs often compromises 

their therapeutic effects; poor specificity of their drug action in general; and the frequent occurrence 

of drug resistance during chemotherapeutic treatment. These issues may all at least be partly 

tackled by delivering anti-tumor compounds with a suitable drug carrier system. SLN is potentially 

a valuable choice for this purpose. The favorable physicochemical characteristics, controlled 

release kinetics, and site-specific drug delivery of SLN mean most of the anti-tumor drugs can 

be efficiently loaded and delivered to the cancerous tissues at reasonable rates while keeping the 

healthy tissues relatively unexposed to these highly toxic agents 17,18 In addition, many compounds 

that can be used to reduce drug resistance in cancer cells are also lipophilic (to be discussed later). 

SLN may also be used for the delivery of these compounds to further improve the effectiveness of 

chemotherapy of cancer normally resistant to cytotoxic drugs. 


Solid Lipid Nanoparticles for Anti-Tumor Drug Delivery 743 

In order to fully optimize the performance of SLN for anti-tumor drug delivery, it is critical to 

gain some understanding of their fundamental aspects. This starts from knowing their main 

constituents that define the physicochemical properties of the SLN system they form, including 

the particle size, surface charge, and lipid crystallinity. These, in turn, have a strong impact on the 

SLN performance. Drug loading, release kinetics, tumor targeting ability, and pharmacokinetics 

in vivo may all be affected. For example, drug encapsulation by SLN is inversely correlated to the 

degree of crystallinity of lipids that is related to the choice of lipids. 19–21 Surfactants are applied to 

stabilize the colloids 22 to delay lipid crystallization 23,24 and modulate the degradation rate of 

lipid. 25,26 The tumor targeting capacity of SLN derives from their particle size, colloidal 

morphology, and surface modification. The leaky vasculature and poor lymphatic drainage of a 

tumor generally facilitate preferential penetration by and retention of particulate matters of small 

size. 27 The submicron size of SLN allows them to passively target tumor tissues. Further manipulation 

of the in vivo distribution of SLN can be achieved by modifying their surface charges or 

hydrophobicity. All in all, by understanding the relationship among the SLN constituents, the 

physicochemical properties, and particle behaviors, the anti-cancer performance of SLN can 

be optimized. 28,29 

SLN are not without their limitations. In fact, some drawbacks of SLN such as the low loading 

capacity for hydrophilic drugs, biphasic release kinetics, and formulation stability issues are still 

barriers against their popular use. 30,31 Fortunately, this class of drug carriers is also rapidly evolving. 

A number of strategies have been developed to overcome these limitations. Some of these 

strategies lead to new generations of SLN. In this chapter, an overall review of the fundamental 

aspects of SLN, the strategies employed for efficient encapsulation and controlled release of various 

anti-cancer drugs by SLN, and their current use for anti-tumor drug delivery is presented. It is 

believed that the information described here will contribute to the design and development of new 

SLN formulations for improved delivery of anti-tumor compounds. 

36.2 COMPOSITION AND STRUCTURE OF SLN 

The typical ingredients of SLN consist of solid lipid(s), surfactant(s) (plus co-surfactants if 

required), and incorporated active ingredients. To date, the lipids applied in the preparation of 

SLN include the following categories based on their structural diversity: fatty acids with different 

lengths of hydrocarbon chain; fatty esters; fatty alcohol; glycerides with different structures; and 

mixtures of glyceryl esters. The frequently used anionic and non-ionic lipids and their structures are 

listed in Table 36.1a and the cationic lipids in Table 36.1b. Waxes such as Cutina CP (cetylpalmitate), 

solid paraffin, and beewax are also sometimes used for SLN preparation. The surfactants 

used include amphoteric, ionic, and non-ionic ones (Table 36.2). Although ionic surfactants such as 

sodium dodecyl sulphate will allow the SLN to possess narrow size distribution and better stability, 

some of them are associated with undesirable toxicity. In addition to the conventional lipid-based 

SLN, the synthesized para-acyl-calix[4]arene-based SLN was also recently developed. 32 

SLN have the inherent advantage to poorly encapsulate water-soluble agents because of the 

lipophilicity of solid lipids. However, to incorporate hydrophilic drugs into SLN, some additional 

measures have to be taken. Counterions such as polymers and esters (organic salts) are required to 

neutralize the charge of the drugs (both listed in Table 36.2). This may increase the apparent 

partition coefficient of the selected drug in the lipid phase and enhance its loading. The encapsulation 

of specific hydrophilic anti-tumor compounds is further discussed in Section 36.5.2.2 and 

Section 36.5.2.3. 

According to the drug incorporation mechanisms, SLN-based systems can be classified into the 

following categories, i.e., classical SLN, polymer–lipid hybrid SLN (PLN), nanostructured lipid 

carriers (NLC), and lipid-drug conjugate (LDC). Their differences are summarized in Table 36.3. 


744 

TABLE 36.1a 

Structures of Lipids Used in the Preparation of Solid Lipid Nanoparticles (SLN) 

C 

H 3 

C 

H 2 

O 

HC O 

C 

H 2 

C H2 

Chemical Structure Name of Lipid 

n 

OC 

OC C H2 

COOH Dodecanoic acid (nZ10) 

Myristic acid (nZ12) 

Palmitic acid (nZ14) 

Stearic acid (nZ16) 

C 

H 2 

C O OC C 

H H2 

2 

O 

HC OH 

C 

H 2 

C 

H 2 

OH 

O 

HC OH 

H 2 

C 

H 2 

C 

2 

OH 

O 

HC OH 

C 

H 2 

C 

H OH 

O 

HC O 

C 

H 2 

H 31 C 15 

OH 

OC C H2 

OC 

C 

H 2 

OC C H2 

OC 

CH 3 

CH 3 

36.2.1 LIPIDS USED IN SLN 

n 

n 

n 

n 

CH 3 

CH 3 

OH 

C C 

n m 

H 

C 

H 2 

OC C H 2 

O C 16 H 33 

CH 3 

H 2 

CH 3 

Caprylate triglycerides (nZ6) 

Caprate triglycerides (nZ8) 

Trilaurin (nZ10) 

Tripalmitin (nZ14) 

Tristearin (nZ16) 

Tribehenin (nZ20) 

Glyceryl monostearate (nZ16) 

Glyceryl hydroxystearate (nZ10, mZ5) 

n Glyceryl behenate (mono-, nZ20) 

n 

CH 3 

n Glyceryl behenate (di-, nZ20) 

CH 3 

O Cetylpalmitate 


The choice of lipids is critical for SLN to achieve the desired drug loading capacity, stability, and 

sustained release behavior. Different lipid phases lead to different apparent partition coefficients 

and different loading capacities for the same drug. Polymorphism of lipids also affects the properties 

of a SLN system. In addition, hydrophobicity of lipids varies with the balance of the 

hydrophobic and hydrophilic functional groups on the lipid molecules. For instance, a fatty acid 



TABLE 36.1b 

Cationic Lipids Used for Preparation of Cationic SLN 

Stearylamine (SA) 

N,N-di-(b-steaorylethyl)-N,N-dimethylammonium chloride (EQ1) 

Benzalkonium chloride (BA) 

Cetrimide (CTAB) 

Dimethyldioctadecylammonium bromide (DDAB) 

N-[1-(2,3-dioleoyloxy)propyl-N,N,N-trimethylammonium chloride (DOTAP) 

Chloroquine phosphate (CQ) 

Cetylpyridinium chloride (CPC) 

Dimethyldioctadecylammonium bromide (DDAB) 

TABLE 36.2 

A List of Surfactants, Co-Surfactants, Counterions, and Surface-Modifying Agents 

Commonly Used in the Formulation of SLN-Based Systems 

Surfactants and Co-Surfactants 

Amphoteric surfactants Non-ionic surfactants 

Egg phosphatidylcholine (EggPC) Brij78 

Egg lecithin (Lipoid E80) Poloxamer 188 

Soy phosphatidylcholine (SP) Poloxamer 407 

(Epikuron 200, 95% SP) Poloxamine 908 

(Lipoid S100) Polyglycerol methyl glucose distearate 

(Lipoid S75, 68% SP) (Tego care 450) 

(Lipoid S75, 68% SP) Solutol HS15 

(Phospholipon 90G, 90%) Span 85 

Trehalose 

Ionic surfactants Tween 20 

Sodium cholate Tween 80 

Sodium cocoamphoacetate Tyloxapol 

Sodium dodecyl sulfate (SDS) 

Sodium glycocholate Co-surfactants 

Sodium oleate Butanol 

Sodium taurocholate Butyric acid 

Sodium taurodeoxycholate 

Counterions 

Organic salts Ionic polymers 

Monodecylphosphate Dextran sulfate sodium salt 

Monohexadecylphosphate Hydrolyzed, polymerized epoxidized soy 

bean oil (HPESO) 

Mono-octylphosphate 

Sodium hexadecylphosphate 

Surface-modifying Agents For long-circulating SLN 

Dipalmitoylphosphatidyl-ethanolamine conjugated with polyethylene glycol 2000 (DPPE-PEG 2000) 

Distearoyl-phosphatidylethanolamine–N-poly(ethylene glycol) 2000 (DSPE–PEG) 

stearic acid-PEG 2000 


746 

TABLE 36.3 

Categories of Solid Lipid Nanoparticles (SLN) 


SLN 1. Solid lipid matrix as carrier 

2. Lipophilic drug or hydrophilic drug–ester complex is molecularly incorporated into 

the solid lipid matrix 

NLC 1. Binary mixture of solid lipid and spatially different liquid oil as carrier 

2. Lipophilic drug or hydrophilic drug–ester complex is molecularly incorporated into 

the solid lipid matrix 

PLN 1. Solid lipid as carrier 

2. Hydrophilic drug–polymer complex is incorporated into the solid lipid matrix 

LDC 1. Hydrophilic drug–lipid conjugate as the homogeneous matrix 

becomes more hydrophobic with the increase in the length of its hydrocarbon chain. A fatty ester is 

more hydrophobic than the fatty acid of the same chain length because the hydrophilic carboxyl 

group of the fatty acid is replaced by the more hydrophobic ester group. Triglycerides are more 

hydrophobic than mono- and di-glycerides because all three hydrophilic hydroxyl groups are 

substituted by fatty ester. The lipid composition essentially determines the overall hydrophobicity 

of SLN. 

Thus far, there has not been a definite rule available for choice of lipids for SLN preparation 

though some empirical guidelines are proposed. Measurement of solubility of a given drug was 

suggested for oral drug delivery by SLN. 33 Empirical equations were used to mathematically 

predict the partition of a drug between a lipid or oil phase and an aqueous phase based on physical 

interactions of the solute (drug) and solvents (lipid or aqueous phase). For parenteral SLN formulations, 

certain requirements need to be met in the selection of lipids such as the suitability of 

producing particulates in a submicron range, sterilization by autoclaving, long-term storage stability 

either in dispersion or after lyophilization, and acceptable toxicity. 34 

In general, high drug loading in SLN can only be obtained when the drug has a high solubility in 

the lipid melt or a high partition coefficient between the molten lipid and aqueous phase. Based on 

the general principle of like-dissolve-like, Li et al. proposed a method to rationally screen the lipid 

carrier for SLN to obtain high drug loading capacity by a combination of the theoretical calculations 

and experimental partition coefficient measurements. 35 The theoretical calculations include 

total solubility parameter, polarity, and mixing enthalpy. This method has already been used for the 

screening of 16 lipids from different categories, and it has shown good predictive power. 

The occurrence of lipid polymorphism is a phenomenon fairly unique to SLN as compared to 

other lipid formulations. Lipid polymorphism is referred to the multiple crystalline forms of solid 

lipids. One of these polymorphic forms, usually the one that forms perfect crystalline lattice, is 

more thermodynamically stable. For example, triglyceride has three forms, i.e., a-form, b-form, 

and b 0 -form. Among them, only the b-form is stable. 36 The relatively less stable or meta-stable 

forms would eventually transform to the stable polymorphic form, leading to a number of challenges 

in the development of SLN formulations. Drug molecules are primarily loaded into SLN in 

the defects of the lattices of solid lipids. When the lipid molecules are converted from the metastable 

forms into the stable form, they become more orderly packed, and some of the defects in the 

lattices disappear. As a result, drug expulsion from the lipid core to the particle surface may occur, 

contributing to high initial burst release and drug leakage during storage. 

Therefore, one additional criterion for choosing lipids for SLN formulation is the tendency of 

these lipids to form perfect crystals, or more precisely, their polymorphic transition rates. Once 

again, there is no strict rule to follow. However, in general, lipids with longer fatty acid chains have 

a slower transition rate than those with shorter chains. 22 SLN made from waxes as lipid matrix 



exhibited significant drug expulsion phenomenon because of the more crystalline structure 

although wax-based SLN are usually more physically stable. 19 In addition to the type of lipids, 

some other factors may also affect the lipid crystallinity, including the storage condition and SLN 

production methods. 37–39 For example, rapid cooling process will be beneficial to maintain the lipid 

matrix in a meta-stable form. 37 The choice of surfactant and co-surfactant plays a crucial role as 

well. 24 The crystallization rate of Dynasan 114-based SLN was shown much faster when poloxamer 

was used as the surfactant than those stabilized by the sodium salt of cholic acid. 40 An increased 

amount of the meta-stable form of stearic acid was also found when a high concentration of 

co-surfactant such as butanol was included in the SLN preparation. 25 

To avoid the aforementioned crystallinity and polymorphism problems, lipid oils, e.g., Miglyol 

812 (caprylic/capric triglycerides) or oleic acid, were incorporated into solid lipids or a binary 

mixture of physically incompatible solid lipids was used, attempting to disrupt the crystallinity of 

the solid lipid matrix. The resulting modified form of SLN is often referred to as nanostructured 

lipid carrier (NLC) (see more description in Section 36.2.4.2). 41–43 This method may also provide 

additional advantages such as increased drug payload. 

Rather than avoiding the complexity caused by lipid polymorphism and crystallinity, some 

researchers attempted to exploit this phenomenon, and they turned it into a useful feature. For 

instance, the polymorphic conversion of lipid matrix from the meta-stable b 0 form to the stable b 

form, as confirmed by form wide-angle x-ray scattering pattern, altered drug release rate, 44 

suggesting that by careful design, release rate of SLN can be modulated by controlling the 

polymorphic interconversion. 

Cationic SLN formulations containing positively charged lipids were developed for gene 

delivery 45 as the positive surface charge may enhance in vivo transfection efficiency of genes. 

Some of the catioic lipids are listed in Table 36.1b. From the point of view of cytotoxicity, twotailed 

cationic lipids are usually preferable to the one-tailed cationic lipids (i.e., CPC and CTAB) 

and cationic surfactants. 

36.2.2 SURFACTANTS USED IN SLN 

The main functions of surfactants in SLN are to disperse the melt lipid into aqueous phase in the 

SLN preparation process and then to stabilize the nanoparticles after cooling. A surfactant molecule 

has a hydrophilic head and lipophilic tail to reduce the surface tension between two phases. The 

hydrophile–lipophile balance (HLB) value of surfactants represents the relative proportion of the 

hydrophilic and lipophilic parts of the molecule. 

As illustrated in Table 36.2, surfactants applied in SLN include four categories: anionic, e.g., 

sodium stearate and sodium dodecyl sulfate; cationic, dodecyltrimethyl ammonium bromide 

(DTAB); non-ionic, e.g., members of the Pluronics w or Tweens w family; and amphoteric, e.g., 

egg lecithin. In general, ionic surfactants employ the electrostatic approach, and non-ionic surfactants 

rely on steric repulsion to stabilize particles. Among them, most non-ionic surfactants consist 

of a hydrophobic moiety (i.e., hydrocarbon chain) and a hydrophilic part (i.e., ethylene oxide 

group). Amphoteric surfactants possess both positively and negatively charged groups and 

exhibit cationic surfactant’s feature at low pH and anionic at high pH. 

Several factors should be considered when choosing surfactants for SLN preparation, namely 

HLB value, route of administration, toxicity, and the effect on lipid modification and particle size. 

Surfactants with HLB values in the range of 8–18 are suitable for the preparation of oil-in-water 

dispersion. Non-ionic surfactants are preferred for oral and parenteral preparations over ionic ones 

because of their lower toxicity and irritancy. In general, the order of the toxicity of surfactants is: 

cationic O anionic O non-ionic O amphoteric. 

One important factor in the choice of a surfactant for SLN should be considered: different 

surfactants have variable influences on the in vivo biodegradation of the lipid matrix. For instance, 

non-ionic surfactants are more effective for inhibiting the degradation of lipid matrix in vivo. 


748 

Poly(ethylene oxide) (PEO) chains of non-ionic surfactants provided steric hindrance of the 

anchoring of lipase/co-lipase complex. Therefore, in vivo degradation rate of SLN may be tailored 

by adjusting the density of PEO chains on the particle surface. For SLN made of the same lipid but 

different non-ionic surfactants, the inhibitory activity against lipid degradation is in the order: 

cholic acid, sodium salt (NaCh) ! lipoid E80 ! Tween w 80 ! poloxamer 407. 25,40 Another 

study shows that Dynasan 114-based SLN have the fastest degradation rate when stabilized by 

lecithin and NaCh; intermediate rate using Tween w 80; and the slowest rate using poloxamer 407. 26 

Mixing NaCh and poloxamer does not affect the degradation rate. 

Finally, the choice of surfactants and co-surfactants also affects the particle size of SLN. SLN 

made of the same lipids may have different sizes because of the use of different surfactants. 25 Zhang et 

al. prepared two types of long-circulating SLN using Brij w 78 or Pluronic w F68 as the surfactant, and 

they found that different surfactants led to the SLNs with totally different particle size. 46 Bocca et al. 

showed that the average diameters of SLN systems produced by using glycocholate as the co-surfactant 

were larger than those using taurocholate. 47 The authors attributed this phenomenon to the 

difference in pKa values of the co-surfactants, i.e., 4.4 for glycocholate and 1.4 for taurocholate. 

36.2.3 OTHER AGENTS USED IN SLN 

In addition to lipids and surfactants (and sometimes co-surfactants), in some modified forms of 

SLN, other agents may also be used. For example, for the encapsulation of cationic, water-soluble 

compounds, a source of counterions is needed that includes organic anions and anionic polymers 

(see Table 36.2). 

SLN can also be surface-modified. When a drug carrier was coated with hydrophilic polymers 

such as poloxamers, poloxamines, or poly(ethylene glycol) (PEG), this so-called stealth or longcirculating 

carrier may reduce the capture by reticuloendothelial system (RES), and it stays longer 

in systemic circulation. 48,49 For a detailed description of the stealth technique, refer to other 

chapters in the book that deal with long-circulating nanoplaforms. This technique has been 

equally valuable to minimize clearance of SLN by phagocytosis. 47 Because of the high surface 

hydrophobicity of SLN, the hydrophilic coating agents are usually pre-conjugated to lipid moieties 

to form amphiphilic molecules (see Table 36.2) so their attachment on to the particle surface can be 

more secure. For the effects of SLN surface modification on the pharmacokinetics and biodistribution 

of anti-tumor agents, refer to Section 36.5.6. 

36.2.4 TYPES, PROPOSED STRUCTURE, AND MORPHOLOGY OF SLN 

The structure and morphology of SLN are affected by several factors such as the physicochemical 

characteristics of the SLN ingredients and the production method adopted. To date, there has been 

no concrete data revealing the detailed structure of SLN. Most SLN researchers believe that SLN 

are probably spherical because a spherical subject has the smallest possible surface area-to-volume 

ratio. This may minimize the surface tension between the lipids and the aqueous phase, thereby 

leading to thermodynamic stability. Nevertheless, it should be noted that the platelet shape of SLN 

and NLC was also claimed by some groups, and they have provided evidence as well. 50,51 

36.2.4.1 Classical SLN 


Based on the difference of melting points (MP) between drug and lipid matrix together with the 

consideration of the release kinetics of SLN, Mühlen et al. proposed three hypothetical models for 

classical SLN: drug-enriched core, drug-enriched shell, and solid solution 52 as illustrated in 

Figure 36.1. Atomic force microscopy (AFM) and DSC studies of prednisolone-loaded SLN 

showed that the particles had an inner crystalline core (hard) and an outside amorphous layer 

(soft). 53 The core-enriched model is based on the fact that the compound with a higher MP will 

precipitate first during cooling. In contrast, if the drug has a lower MP than the lipid matrix, 



(a) (b) 

(d) (e) 

FIGURE 36.1 Proposed structures of solid lipid nanoparticles (SLN): (a) drug-enriched core model; (b) drugenriched 

shell model; and (c) drug molecularly dispersed model. (Adopted from zur Mühlen, A., Schwarz, C., 

and Mehnert, W., Eur. J. Pharm. Biopharm., 45, 149–155, 1998.); and nanostructured lipid carrier (NLC): (d) 

micro-compartment model; (e) liquid oil molecularly dispersed model. 

a drug-enriched shell model will fit the particles. If the drug has a similar MP to the lipid matrix, a 

solid solution model will best describe the particle. Although there is no formal study on the 

structure of LDC, it should be similar to classical SLN with the loaded drug evenly distributed 

within the lipid matrix because they are directly bonded to the lipid molecules. 

36.2.4.2 Nanostructured Lipid Carrier (NLC) 

To overcome the difficulties of classical SLN, NLC was introduced. In this case, the lipid matrix is 

composed of the binary mixture of a solid lipid and the medium chain triglycerides or liquid oil. 

BasedonthecompositionofthelipidmatrixofNLC,Jenningetal.proposedtwodrug 

incorporation models 54 as demonstrated in Figure 36.1. In the first model, liquid oil is molecularly 

dispersed within the solid lipid matrix when the concentration of the liquid oil is below its solubility 

in the lipid. In the second model, liquid oil is distributed in the solid lipid in droplet form. Physically 

separate phases (solid, liquid) coexist in this model. 

36.2.4.3 Polymer–Lipid Hybrid Nanoparticles (PLN) 

In order to deliver the hydrophilic drugs with a high drug loading capacity and simultaneously 

control the release kinetics of SLN, a new variation of SLN, PLN was recently developed. 55,56 The 

schematic illustration of proposed the formation mechanism and structure of PLN is presented in 

Figure 36.2. An ionic drug, e.g., doxorubicin HCl, forms complex with a counter ionic polymer, 

e.g., dextran sulfate that then partitions into the lipid phase because of higher hydrophobicity, 

enabling incorporation of water soluble drugs in the lipid nanoparticles. PLN have a spherical 

Water-soluble 

counter polymer 

+ 

Water-soluble 

ionic drug 

Drug-polymer 

complex 

(c) 

+ lipid 

Drug-loaded polymer-lipid 

Hybrid nanoparticles (PLN) 

FIGURE 36.2 Proposed formation mechanism and structure of polymer–lipid hybrid nanoparticles (PLN). 


750 

FIGURE 36.3 Transmission electron microscope (TEM) image of polymer–lipid hybrid nanoparticles (PLN) 

loaded with doxorubicin. (From Wong et al., Pharm. Res., in press.) 

morphology as shown in Figure 36.3. Modest initial burst release and subsequent sustained release 

profiles have been observed using this system. However, if the drug–polymer complex is uniformed 

distributed in the solid lipid matrix as shown in Figure 36.2 or just attached to the surface of 

particles needs further verification. 

Li et al. applied differential scanning calorimetry (DSC) and powder x-ray diffraction (PXRD) 

crystallography to examine crystallinity and crystal structures of PLN made of verapamil HCl, 

dextran sulfate, and dodecanoic acid. 35 The DSC and the PXRD results (Figure 36.4) showed that 

Intensity 

2 10 20 30 

2 theta, degrees 

40 50 

FIGURE 36.4 Overlaid powder x-ray diffraction crystallographs of (1) verapamil HCl (VRP); (2) dodecanoic 

acid; (3) dextran sulfate sodium salt (DS); (4) physical mixture of verapamil HCl and dextran sulfate sodium 

salt; (5) VRP–DS complex; and (6) VRP–DS complex incorporated dodecanoic acid. (From Li et al., Pharm. 

Res., 2005.) 



6 

5 

4 

3 

2 

1


the crystal structure of the lipid remained, but its melting temperature decreased a bit as the drug 

became amorphous in the complex form. This finding suggests that the presence of drug–polymer 

complex in PLN may be similar to the liquid oils that reduce the lipid crystallinity in NLC. 

Nevertheless, the exact significance of the finding will require further investigations. 

36.3 MANUFACTURING METHODS 

Currently, many techniques are available for SLN preparation. These include high pressure homogenization, 

57 microemulsion, 58 and solvent diffusion. 59,60 The more commonly used methods will 

be briefly described in this chapter. For detailed descriptions of SLN preparation, please refer to the 

up-to-date reviews 30,38,57–60 where the optimization of the processing variables is also discussed. 

36.3.1 HIGH PRESSURE HOMOGENIZATION 

High pressure homogenization methods include hot and cold approaches. The main advantages of 

these methods include narrow size distribution of nanoparticle produced and the possibility of 

scale-up production. The cold approach minimizes the exposure of drug to heat (although brief 

period of heating is still required) and is more suitable for labile drugs sensitive to heating. 

However, this approach requires more tedious steps, and it results in particles of more variable 

sizes. The general scheme of high pressure homogenization method is shown in Figure 36.5 (hot 

method) and Figure 36.6 (cold method). 

36.3.2 MICROEMULSION 

A warm microemulsion is prepared under stirring before it is dispersed in a large volume of cold 

water at 2–48C. The typical volume ratio of emulsion to cold water is in the range of 1–20. The 

procedure of microemulsion method for SLN preparation is illustrated in Figure 36.5. 

Hot homogenization 

method 

High-speed stirring to 

form coarse emulsion 

High pressure homogenization 

at a temperature above the M.P. 

of lipid 

Cool down 

Melt lipid and 

disperse drug in lipid 

Mix the molten lipid and 

drug with hot surfactant 

solution 

Solid lipid nanoparticles 

Microemulsion 

method 

Gentle stirring, high surfactant/ 

cosurfactant concentration leads 

to microemulsion 

Dispersing into excess 

cold water under stirring 

FIGURE 36.5 Schematic procedures of hot homogenization and microemulsion for SLN production. 


752 

Cold homogenization method 

Melt lipid, disperse 

drug in lipid melt 

Cool to solidify lipid/drug mixture 

Ground into fine powder 

under liquid nitrogen 

Dispering the lipid/drug powder 

into a cold surfactant solution 

High-speed stirring to 

form coarse emulsion 

High pressure homogenization at 

low temperature 


FIGURE 36.6 Schematic procedure of cold homogenization for SLN production. 

Using this method, it is noted that the particle size, size distribution, and stability of SLN is 

correlated with the properties of microemulsion such as the lipids and surfactants applied, the 

concentration of surfactant, and the drug/lipid ratio. 

Solvent evaporation method 

Lipid and drug dissolved in water-immiscible 

organic solvents, i.e., cyclohexane or toluene 

Mix with the aqueous surfactant 

solution 

High-speed stirring to produce 

coarse emulsion 

High pressure homogenization 

Vacuum evaporation of organic solvents 


FIGURE 36.7 Schematic procedure of solvent evaporation method for SLN production. 




36.3.3 SOLVENT EVAPORATION 

As demonstrated in Figure 36.7, the lipid is dissolved in water-immiscible organic solvent first, and 

then the hydrophobic drug is added. Dispersing the drug–lipid–organic solvent mixture into an 

aqueous solution of emulsifiers produces an o/w emulsion. SLN are obtained by removing the 

organic solvent with evaporation under low pressure. 

This method totally avoids heating. Its main disadvantage is the inclusion of organic solvent in 

preparation. Residues of organic solvent may raise the bio-safety concern. 

36.3.4 W/O/W DOUBLE EMULSION METHOD 

This method is designated for the delivery of hydrophilic drugs by using SLN. 61 Highly concentrated 

drug solution is dispersed into a lipid phase under vigorous stirring and stabilized by a 

surfactant to form the primary w/o emulsion. The primary emulsion is subsequently dispersed 

into a large volume of aqueous solution containing another kind of surfactant to form w/o/w 

emulsion. The nanoparticles are obtained by ultrafiltration or solvent evaporation. The detailed 

procedure is illustration in Figure 36.8. 

36.4 CHARACTERISTIC PROPERTIES OF SLN 

36.4.1 DRUG LOADING 

Double emulsion method 

Dispersing small amount of 

hydrophilic drug solution 

in molten lipid 

Adding stabilizers 

Primary w/o emulsion 

Dispersing into a large volume of 

aqueous phase containing stabilizers 

under higher speed stirring 

Cooling down 


FIGURE 36.8 Schematic procedure of w/o/w double emulsion method for SLN production. 

For classical SLN, NLC, and PLN, the drug loading process involves partitioning and re-partitioning 

of the drug between the lipid and aqueous phase during the heating and subsequent cooling 

process. No covalent bonding or ionic bonding between the molecules of the drug and lipid occurs. 

On the other hand, drug incorporation by LDC, another modified form of SLN, involves direct 

covalent or ionic bonding of the drug molecules to the lipids. 

Two models are currently proposed to describe how drug molecules are incorporated in SLN, 

NLC, and PLN. In the first model, drug molecules are accommodated in the flaws of the lipid crystal 

lattice. This model has been widely used to interpret the drug loading behaviors and the biphase 


754 

release kinetics of SLN and NLC. In the second model, the physical and chemical compatibilities 

(affinity, miscibility) between the drug–polymer complex and the solid lipid phase are used to 

predict the drug loading by PLN. 35 

For water-soluble, ionic anti-tumor compounds, to allow their efficient encapsulation by SLN, 

three possible strategies have so far been employed. These strategies include mixing the anti-tumor 

compound with an organic salt carrying opposite charges and packaging the resulting ion-pairs into 

lipids; 62–65 complexing the ionic anti-tumor compound to lipids carrying opposite charges and 

directly using the drug–lipid complexes for SLN preparation (i.e., LDC preparation); and 

complexing the ionic anti-tumor compound to polymer molecules carrying opposite charges, i.e., 

polyelectrolytes and dispersing these drug–polymer complexes into lipids to form SLN (i.e., PLN 

preparation). Figure 36.9 summarizes these three feasible strategies. These strategies are similar in 

terms of their underlying principle—to neutralize the charge on the drug with a counter ion to 

facilitate drug partitioning into the lipids. 

When an organic salt or ionic polymer interacts with the drug molecules of opposite charge, the 

resulting complex loses charges and gains in hydrophobicity that favors their incorporation into the 

solid lipid matrix. The ionic molar ratio between the organic salt or polymer and the drug plays a 

critical role in this process. By employing isothermal titration calorimetry in conjunction with 

partition experiments, Li et al. found that at an equal ionic molar ratio of verapamil HCl (VRP) 

to dextran sulfate (DS), the VRP-DS complex reached maximum hydrophobicity and partition in 

the lipid phase. 35 

36.4.2 DRUG RELEASE 

(1) Drug + + Ion – 

(2) Drug + + Lipid – 

Lipid 

(3) nDrug + + Polymer n– 

Lipid 

Drug + –Lipid – 

Drug-Ion-SLN 


[Drug + –Lipid – ]-SLN 

[nDrug + – Polymer n– ]-SLN 

FIGURE 36.9 Strategies to load cationic, water-soluble anti-tumor compounds into SLN-based systems. (1) 

Ion-pair formation method; (2) lipid drug conjugate (LDC); and (3) polymer–lipid hybrid nanoparticles (PLN). 

The release kinetics of SLN usually exhibit a biphasic feature: a strong initial burst release (usually 

O 50% of the loaded drug in the first several minutes or hours) followed by a subsequent prolonged 

release. It is a consensus that the release behavior of SLN is determined by the physicochemical 

properties of the constituents and the morphology of SLN. By correlating the drug release kinetics 

to the physicochemical properties of the drug and lipids, the hypothetical structures of SLN can be 

deduced (as mentioned in Section 36.2.4.1). In general, sustained release profiles are correlated to 

the drug-enriched core model, and both the drug-enriched shell model and the solid solution model 

will have fast release kinetics. 

High drug loading capacity can be achieved in NLC. However, the existence of the microcompartments 

of oil droplet in the solid lipid matrix may compromise the control over the release 

kinetics. The control of release can be improved in the NLC model with liquid oil molecularly 

dispersed in the solid lipid matrix. 

In comparison to SLN and NLC, PLN offer an additional mechanism to regulate the drug release 

kinetics. In addition to the regular diffusion control, the need for ion exchange between the drug– 

polymer complex and the surrounding medium for the drug to be released may also be exploited for 

better drug release control. The synergistic effect of both approaches will possibly lead to reduced 



% Dox released from Dox-PLN 

100 

initial burst release and a subsequent sustained release. Figure 36.10 presents representative release 

profiles of PLN. Moderate initial releases (50% w/w of drug released in first few hours in both dextran 

sulfate-based or HPESO-based PLN loaded with doxorubicin) were achieved when 0.05M CaCl 2 

was used as the release medium, and sustained releases were observed in the late phase of release. 

36.4.3 STABILITY OF SLN 

90 

80 

70 

60 

50 

40 

30 

20 

10 

0 

0 2 4 6 8 10 12 14 16 

Time (hours) 

Dox-PLN (HPESO) 

Dox-PLN (DS) 

FIGURE 36.10 Cumulative percent of doxorubicin released versus time for PLN in 0.05 M CaCl 2 solution at 

378C. ( ) Dextran sulfate (DS) sodium salt as the counter polymer; and ( ) hydrolyzed polymer of 

expoxidized soybean oil (HPESO). (Adapted from Wong et al., Pharm. Res., in press.) 

The practical usage of SLN depends on their long-term physical and chemical stability. Typically, 

the shelf life of one viable formulation should be at least one year. The criteria for the assessment of 

physical stability of SLN include average particle size, zeta potential, and drug leakage during 

storage. For labile compounds, it is critical to evaluate their chemical stability as well. 

The physicochemical properties of the materials used to prepare SLN such as the lipids, drugs, 

and surfactants play a crucial role in determining the stability of SLN. Waxes usually lead to slower 

particle growth and aggregation than glycerides. 19 Glycerides that include a fraction of monoglycerides 

are usually more stable than those without (e.g., Dynasan 116 O Compritol 888 ATO O 

Imwitor 900). Ionic surfactants often stabilize SLN better than non-ionic surfactants. 

In addition to the physicochemical properties of the SLN material, the stability of SLN is also 

strongly affected by the various processes the SLN subject to after their preparation. For instance, 

sterilization is strictly needed for parenteral SLN formulation. When compared to most other drug 

delivery systems, SLN are surprisingly stable when challenged by the harsh conditions 

(compressed, saturated steam at 1218C) during autoclaving. Cavalli et al. found no significant 

changes occurred for SLN in terms of the particle size, polydispersity (PI) index, and zeta potential 

before and after sterilization. 66 However, there are conflicting reports. Venkateswarlu et al. 39 

demonstrated an increase of SLN size almost 2–3-fold, a shift of zeta potential from positive to 

negative, and a slight drop in drug loading after autoclaving of SLN. The differences between the 

two cases are probably attributed to the different lipids, surfactants, and drugs used in their studies. 

Lyophilization is another process that has a strong influence on the formulation stability. 

Lyophilization, or freeze drying, is necessary to prevent SLN aggregation or hydrolysis in 

aqueous medium. In general, carbohydrate cryoprotectants such as trehalose may stabilize a 

SLN formulation during lyophilization. 67 Proper freezing velocity and re-dispersion method 

(e.g., sonication) are also required. It has been demonstrated that with the optimized parameters 


756 

for lyophilization, the reconstituted SLN are feasible for i.v. injection; 68 although, there are also 

studies showing that some deterioration in the particle quality, particularly the particle size, may 

occur. 69 Even in these studies, the reconstituted SLN remained sufficiently small in size and stable 

for oral administration. 

During storage, the stability of SLN will also be affected by the energy introduced from outside 

sources such as storage temperature, light, and packing materials. 70 Strong light and high temperature 

will accelerate the decrease of the particles’ zeta potential and, in turn, induce particle 

aggregation because colloids with zeta potential less negative than K15 mV are prone to aggregation. 

Some packing materials such as siliconized vials also exhibit statistically significant effects on 

the particle size growth in comparison to that packed in untreated glass vials. 

Most aforementioned stability data derive from non-biological studies. There have been more 

in vivo SLN studies recently on this issue. This is important because the environment of the 

systemic circulation is actually harsh on SLN. For example, the electrolytes in the plasma will 

lead to the shrink of double layer, the rapid fall off of zeta potential, and the subsequent coagulation 

of particles occurs. 71 Particles may also grow in size and subsequently form semi-solid gels through 

adding mono-, di-, and trivalent ions. 72 In addition, the biodegradation of the lipid matrix caused by 

enzymes in vivo will accelerate drug release from SLN and expose the agents to the various 

degradative systems before they can reach the disease site. In brief, even when performing SLN 

studies in vitro, it will be advisable to at least use media that simulate the composition of body fluids 

such as Hanker’s buffer salt solution (HBSS) for the experimental works. This may help the 

researchers to gain insight into the behaviors of SLN in vivo. 

Finally, it must be re-emphasized that the various properties of SLN are strongly affected by the 

occurrence of lipid polymorphism, and these include stability as well. During storage of SLN, drug 

expulsion and subsequent drug leakage may occur as a result of lipid polymorphic transition, 

causing re-distribution of drug molecules within particles. For details of polymorphism, refer to 

Section 36.2.1. All in all, the factors that accelerate particle size growth as a result of the loss of 

surface charge, including high storage temperature and strong light, will also speed up polymorphic 

transition rate. 

36.5 SLN FOR ANTI-TUMOR DRUG DELIVERY 

36.5.1 CANCER CHEMOTHERAPY 


Although hormonal therapy and immunotherapy are useful for certain types of cancer, chemotherapy 

with cytotoxic drugs remains the most established and commonly used form of drug therapy. Sometimes, 

chemotherapy with cytotoxic drugs is the only treatment a patient receives. More often, 

cytotoxic drugs are used in addition to other modalities (e.g., surgery, radiotherapy) to improve 

their effectiveness and prevent cancer recurrence. 73 Until recently, most of the SLN systems for 

cancer treatment are designed for cytotoxic drug delivery. Therefore, unless otherwise specified, in 

the following discussions the terms cytotoxic drug, anti-cancer drug and anti-tumor drug will be 

interchangeably used, even though the latter two terms actually carry broader meaning. 

Cytotoxic drugs treat cancers by preferentially causing death or arrest in growth of cancer cells. 

However, their cytotoxicity is not highly specific for cancer cells. The healthy cells, especially 

those rapidly dividing, almost inevitably suffer from the chemotherapeutic treatment as well. 74 The 

major goal in cancer chemotherapy is, therefore, to achieve reasonably high drug concentration for 

effective cancer cell kill without causing strong toxicity. In addition, clinical resistance to 

chemotherapy drug therapy often occurs in cancer management. 75 One prominent form of drug 

resistance is related to the expression of ATP-dependent membrane-associated drug transporters in 

cancer cells. Cells with over expression of these drug transporters are resistant to a broad range of 

structurally diverse anti-cancer compounds, a phenotype often referred to as multi-drug resistance 



(MDR). 75,76 Cytotoxic drugs can be actively transported out of the cancer cells by these transporters, 

rendering the chemotherapy ineffective even when an adequate dose of drug is administered. 

36.5.2 DELIVERY OF CYTOTOXIC COMPOUNDS USING SLN 

SLN have been quickly gaining ground as a carrier of anti-tumor drugs. 10,11,55,62,77–90 Some of 

these SLN systems have been evaluated for their anti-tumor activities, stabilities, and/or biodistribution 

in biological systems such as cultured cell lines and animal models. These systems and 

TABLE 36.4 

SLN Formulations Used for Delivery of Drugs with Anti-Cancer Properties and a Brief 

Summary of Their Works 

Drugs Groups a 

F/C b 

Focus of Studies 

In Vitro In Vivo P/B c 

Stealth Comments/References 

Camptotecan Yang Y 

Poorly Water-Soluble 

Y(mice) Stability study/ 91,92 

Cholesteryl Serpe 

butyrate 

73 

Y In colon cancer cell 

line/ 73 

Etoposide Murthy Y Y Routes of admin 

compared/ 93 

Irinotecan Unger Y Y Y SN38, irinotecan prodrug 

used, stability 

studied/ 94,95 

Paclitaxel Gasco Y 

Lee Y Y Y Paclitaxel prodrug 

used/ 97 

Muller Y 

98 

Serpe Y In HT28 colon cancer 

cell line/ 73 

Zhang Y Y(mice) 2 formulations using 

different surfactants 

compared/ 99 

Retinoic acid Kim Y In various human cancer 

cell lines/ 90 

Idarubicin Gasco Y 

Ionic, water-soluble 

Y (rabbit) Duodenal route/ 101 

Doxorubicin Gasco Y Y (rat), 

(rabbit) 

Y 

8,100,102 

Serpe Y In colon cancer cell 

line/ 73 

Wu Y Y In MDR breast cancer 

cells, cytotoxicchemosensitizer 

combination/ 10,60 

5-Flurouracil Wang Y 

Non-ionic, water-soluble 

Y(mice) FudR, derivative of 

5-fluouracil used/ 103 

a 

According to corresponding author. 

b 

F/C- formulation and characterization. 

c 

P/B Pharmacokinetics and/or biodistribution studies presented—pharmacokinetics. 


758 

related work are listed in Table 36.4. To this point, studies of SLN systems have demonstrated the 

following properties of SLN that are desirable for anti-tumor drug delivery: 

1. Versatility and modifiability of the carrier that allow encapsulation of cytotoxic agents of 

vastly diverse physicochemical properties; 

2. Improved anti-cancer drug stability; 

3. Capability to load more than one drug in one carrier system that opens up the possibility 

of a drug combination delivery (e.g., a cytotoxic agent and a chemosensitizer); 

4. Improved in vitro cytotoxicity against cancer cells, including those normally refractory 

to chemotherapy; 

5. Enhanced drug efficacy in animal models; and 

6. Improved pharmacokinetics and in vivo drug distribution. 

In the following sections, the use of properties (1) and (2) by SLN for lipophilic anti-tumor drug 

delivery will be jointly discussed in Section 36.5.2.1, and Section 36.5.2.2 through 36.5.2.5 will 

focus on properties (3) to (6), respectively. 

Anti-tumor drugs can be categorized based on their solubility properties. In general, the 

majority of anti-tumor agents, particularly the cytotoxic compounds, are fairly lipophilic and 

poorly water-soluble. 91 There are several exceptions. Some anti-tumor drugs are water-soluble 

because their molecules are small and contain multiple hydrophilic groups. Examples of these 

water-soluble drugs include 5-fluorouracil (5-FU) (Figure 36.11a, MW Z130.1) and mitomycin-C 

(MW Z334.3). In addition, there are some lipophilic anti-tumor molecules that can be 

converted into water-soluble, ionic salts because of their ionizable basic functional groups. These 

salts are much more commonly used in clinical settings than their corresponding water-insoluble 

free bases because they can be conveniently dissolved and administered using standard aqueous 

diluents such as 5% dextrose and 0.9% saline. For example, the hydrochloride salts of anthracyclines, 

e.g., doxorubicin and idarubicin (Figure 36.11b), and sulfates of alkaloids, e.g., vincristine 

and vinblastine, are all commercially available and frequently used in cancer chemotherapy. 

As previously discussed, the efficiency and extent of drug incorporation into SLN are strongly 

influenced by the partition behavior of the drug in a lipid–water two-phase system. The lipophilic, 

uncharged anti-tumor drugs may be encapsulated with relative ease using classical SLN preparation 

methods. The other two classes (uncharged, hydrophilic compounds and charged, lipophilic 

compounds) of water-soluble anti-tumor agents, on the other hand, require new strategies to 

enhance their incorporation into SLN. Means to control the release rates of these compounds 

F 

O 

N 

H 

N 

H 

O 

(a) 5-fluorouracil 

R1 

O 

O 

OH 

OH 

O 

R2 

OH 

H 

O O CH3 *NH2 (b) Doxorubicin /Idarubicin 


FIGURE 36.11 Examples of the molecular structures of two different types of water-soluble anti-cancer 

compounds. (a) 5-fluorouracil (5-FU); and (b) doxorubicin (R 1ZOCH 3,R 2ZCH 2OH) or idarubicin (R 1Z 

H, R 2ZCH 3). * Basic amine group that is protonated when forming the water-soluble hydrochloride salt of 

doxorubicin or idarubicin. 


OH


from SLN are also necessary because they may rapidly diffuse from the drug carriers into the 

surrounding aqueous medium. 

36.5.3 POORLY WATER-SOLUBLE COMPOUNDS 

36.5.3.1 General Principle 

There have been several SLN systems successfully formulated for encapsulation of poorly watersoluble 

anti-tumor compounds. These include camptothecin-based topoisomerase I inhibitors, 

including camptothecin and irinotecan and their prodrugs; taxanes, mainly paclitaxel; and other 

lipophilic compounds with anti-tumor properties. In general, these poorly water-soluble 

compounds partition reasonably well in the lipid phase and can be efficiently encapsulated by 

the conventional SLN without much nanoparticle modifications required. These agents, however, 

carry other problems such as poor stability and likelihood of precipitation. In fact, each group of the 

compounds has its specific problems to be solved before they can be properly delivered using SLN. 

36.5.3.2 Topoisomerase I Inhibitors (Camptothecin, Irinothecan) 

Camptothecin is a plant alkaloidal compound extracted from Camptotheca acuminata. 92 It serves as 

the prototype compound of a relatively new class of anti-tumor agents known as topoisomerase I 

inhibitors (e.g., irinotecan, topotecan). The molecules of camptothecin and its related compounds 

all possess extensive aromatic structures and are quite lipophilic (Figure 36.12a). 93 The anti-tumor 

activity of camptothecin is strongly correlated to the functionality of the lactone ring of the drug 

molecule. However, in an aqueous environment, particularly at basic pH, this lactone ring is 

vulnerable to hydrolysis, leading to the formation of therapeutically inactive carboxylate. The 

poor water solubility and chemical stability of this drug has seriously diminished its practical 

value. These two shortcomings can both be overcome by the use of SLN for their capabilities to 

efficiently load lipophilic compounds and provide a lipid environment to prevent hydrolytic 

degradation. Yang et al. 79 described a SLN system for camptothecin delivery. A very high encapsulation 

efficiency (99.6%) of camptothecin was achieved. The same study also showed that the 

drug mostly remained in its active lactone form until it was released, confirming the protective 

effect of SLN against hydrolytic degradation. 

Irinotecan (or Camptosar or CPT-11) is a newer member in the topoisomerase I inhibitor 

family. It is relatively hydrophilic and stable compared to camptothecin and has been put to clinical 

use for colon cancer chemotherapy. However, compared to its parent drug SN-38 (Figure 36.12b), 

the anti-tumor activity of irinotecan is nearly a thousand-fold weaker. 94 SN-38 is prevented from 

being put to clinical use mainly because of its extreme hydrophobicity. Moreover, hydrolytic 

opening of the labile lactone ring of SN-38 and irinotecan is still possible, though less likely. To 

N 

N 

O 

N 

N 

O 

O 

H3C OH 

O 

– 

O 

OH 

HO 

OH 

H3C OH O 

(a) (b) 

– 

CH 2CH 3 

N 

H 3 C 

N 

O 

O 

O 

OH 

FIGURE 36.12 Molecular structures of two topoisomerase I inhibitors. (a) The hydrolytic degradation of 

camptothecin from its active lactone form (left) to its inactive carboxylate form (right). (b) SN-38, a more 

potent lipophilic parent drug of irinotecan. 


760 

improve its water solubility and chemical stability, SLN formulation of SN-38 was prepared. 81 The 

concentrations of SN-38 in the SLN formulation reached as high as 1.0 mg/ml. The lactone ring 

stability of the compound was also improved. After 3 h incubation in human serum albumin, the 

lactone concentration in SN-38 loaded-SLN remained 80%. This was superior to the retentions of 

only 40% lactone for unformulated SN-38 and irinotecan and even more so when compared to free 

camptothecin that lost more than 95% of its active lactone form in 90 min. 

In brief, SLN are a useful drug carrier for camptothecin-based anti-tumor compounds. Both of 

the studies mentioned above have demonstrated the value of SLN for drug solubilization and 

preservation of drug stability. 

36.5.3.3 Paclitaxel 


Paclitaxel (Taxol) is an anti-microtubule agent with broad spectrum anti-tumor activities. It is 

gaining popularity among oncologists for its effectiveness against several types of malignancies, 

even when used alone. 95 Paclitaxel is a poorly water-soluble drug. At present, it is commercially 

available as a non-aqueous micellar solution containing a polyoxyethylated castor oil Cremophor 

EL and 49.7% dehydrated ethanol. Cremophor EL is known to cause serious hypersensitivity 

reactions and nephrotoxicity in human subjects. 96 In order to solubilize paclitaxel without the 

need for Cremophor EL, a few SLN formulations have been developed and studied. 62,82–86 In 

general, the drug loading in these SLN formulations ranged from 2 to 5%. In these SLN formulations, 

surfactants or stabilizers such as Pluronic w F68 (i.e., poloxamer 188), Brij-78, and 

phosphatidylcholine were used. In comparison to Cremophor EL, all of these excipients are 

commonly included in parenteral formulations with better track records of safety. 

Another major issue associated with the conventional paclitaxel formulation is its poor physical 

stability. Paclitaxel is routinely diluted to a concentration of 0.6–1.2 mg/ml with saline or dextrose 

for clinical use. The dilution can destabilize the formulation and drug precipitation frequently 

occurs, necessitating filtration of the drug solution prior to administration. This extra step adds 

to the labor cost and may affect the drug concentration. This problem is apparently not found in 

paclitaxel-SLN formulations. In one of these studies, 62 the SLN formulation of paclitaxel was even 

diluted to nanomolar range using cell culture medium for in vitro cytotoxicity evaluation. There 

was no report of drug precipitation. 

Drug precipitation also happens when paclitaxel is too quickly discharged from a carrier system 

and the concentration of the released drug exceeds the saturation point. One solution is to limit the 

rate of drug release. As demonstrated in the paclitaxel-loaded SLN system by Cavelli et al., 82 no 

drug precipitation was observed because only 0.1% of paclitaxel was released in pseudo-first order 

in 2 h. 

Besides SLN, researchers have also successfully encapsulated paclitaxel in other drug delivery 

systems, particularly those that are lipid emulsion based. 97–99 However, all of these formulations 

are prone to drug precipitation during storage. It is believed that paclitaxel has limited solubility in 

lipid as well, and the limited solubilities of paclitaxel in both aqueous and lipid phases leads to the 

strong tendency of drug precipitation when the formulations are stored for a long term. Stevens 

et al. 84 described the use of a lipophilic paclitaxel prodrug (paclitaxel-2 0 -carbonyl-cholesterol) to 

overcome this difficulty. Despite the strong lipophilicity of the prodrug, this group was able to 

demonstrate over 10% of drug release into bovine serum albumin in a relatively linear fashion 

within 2 h. In addition, even without resorting to the use of prodrug, Cavalli et al. showed that their 

formulation containing the unaltered paclitaxel was autoclavable at 1218C without a significant 

reduction in the amount of the incorporated drug. 82 The same formulation was also stable for over 

18 months after lyophilization, and the lyophilized nanoparticles appeared easily dispersable when 

cryoprotectant (2% trehalose) was added. 

Therefore, the main shortcomings of paclitaxel, including a high risk of hypersensitivity and 

frequent drug precipitation, are correctable using SLN. A number of new, more cancer-type- 



specific paclitaxel analogs have been synthesized in the past decade. 100,101 It will be interesting to 

see if SLN are equally useful for these compounds. 

36.5.3.4 Other Poorly Water-Soluble Anti-Tumor Compounds (Etoposide, 

All-Trans Retinoic Acid, Butyric Acid) 

Etoposide, all-trans retinoic acid (ATRA), and butyric acid are all lipophilic molecules that have 

anti-tumor properties. Out of these three compounds, etoposide (Vepesid, VP-16) has the longest 

history in cancer chemotherapy and is still in clinical use for the treatment of several forms of 

cancer. Etoposide is a plant alkaloid. It inhibits topoisomerase II and produces reactive derivatives 

that can lead to DNA strand break. 102 This drug was successfully encapsulated in SLN made of 

tripalmitin. 80 The authors quoted that the particles were stable after four months of preparation with 

excellent redispersibility. However, probably because the focus of that study was on the evaluation 

of drug biodistribution and in vivo anti-tumor efficacy, few details in this area was revealed. The 

SLN system demonstrated an excellent encapsulation efficiency of 98.96% with 4% payload of 

etoposide obtained. 

Both ATRA and butyric acid remained at the experimental stage in terms of cancer therapy. 

ATRA is well documented for its chemical instability. 103 It is sensitive to heat, light, and oxidation, 

and it may isomerize into isotretinoin or be oxidized to form inactive products such as all-trans-4oxo. 

This compound should be well incorporated into SLN for its high lipophilicity; although, in the 

only SLN system documented, the drug payload attempted was merely 2.5 mg ATRA per g of SLN 

powder. 77 The chemical stability of ATRA in SLN during storage was evaluated. More than 90% of 

the encapsulated drug remained intact after one month of storage at 48C versus less than 60% when 

the drug was stored in the forms of methanol sol ution or Tween 80 solution under the same 

conditions. Butyric acid is a short chain fatty acid that shows in vitro inhibitory effect against 

colon cancer cell growth. However, its in vivo activity is very low because of its rapid metabolism. 

99 It was believed that by using its prodrug cholesteryl butyrate and by delivering this prodrug 

using SLN, the anti-tumor activity could be preserved. The prodrug was successfully loaded in 

SLN, and the prodrug demonstrated good in vitro activity. 62,104,105 Although the in vivo activity has 

not yet been confirmed, the studies provided another example of the value of SLN to protect a 

labile chemical. 

36.5.4 WATER-SOLUBLE IONIC SALTS 

36.5.4.1 General Principle 

The anti-tumor drug salts typically have lipophilic molecular structures. The main obstacle against 

these salts from efficient loading into SLN is their ionic charges. For details of the drug loading 

mechanism of this class of compounds into SLN-based systems, refer to Section 36.4.1. In essence, 

it is about neutralizing the charge on the ionic drug salt with a counter ion. It may be argued that 

these extra steps can simply be avoided by using the free bases of these agents. However, the use of 

free base compounds will essentially lead to the same scenario described previously in Section 

36.5.1.1. These poorly water-soluble free base compounds will likely be very slowly released, and 

the released drug concentrations will not be as high as compared to when the salt forms are used. 

The free base compounds may also require dissolution in organic solvents first during the preparation 

process to allow even mixing with the lipids. In general, it is preferable to use the watersoluble 

salt of an anti-tumor drug for SLN formulation whenever it is available. 

36.5.4.2 Doxorubicin and Idarubicin 

Although there are quite a few anti-tumor drugs available in ionic salt forms for improved water 

solubility, only two of the anthracyclines have been studied for delivery by SLN-based 


762 

formulations. These include doxorubicin and idarubicin. These two compounds are similar in 

structures except that doxorubicin is more hydrophilic because of the presence of additional 

polar substituent groups (Figure 36.11b). Each of these two drug molecules carry an amine 

group and can be converted into hydrochloride salts. To facilitate the loading of these salts into 

SLN, they need to be made lipid-soluble based on one of the above-mentioned strategies. Currently, 

there is still no LDC system designed for anti-tumor drug delivery. The other two strategies 

described in Section 36.4.1 (ion-pair, PLN) have both been used and tested. 

Gasco’s group used decyl phosphate or hexadecyl phosphate to form ion pairs with doxorubicin 

hydrochloride and idarubicin hydrochloride to enhance their loading of into SLN made of stearic 

acid and egg lecithin. 87 The ion pair formation resulted in increases in lipophilicity as defined by 

apparent partition coefficients between water and stearic acid at 708C by more than 1000-fold for 

doxorubicin and more than 300-fold for idarubicin. SLN prepared using these lipophilic ion pairs 

carried payloads of doxorubicin and idarubicin up to 7% and 8.4%, respectively. Probably because 

of the high lipophilicity, the drug also released very slowly. For both drugs, less than 0.1% drug was 

released after 2 h. With these release rates, there lies a theoretical risk that only the cancer cells in 

physical contacts with the SLN may be exposed to sufficiently high drug concentrations and 

properly treated. In addition, as demonstrated in several studies, chronic exposure to sub-lethal 

concentrations of cytotoxic drugs can induce expression of Pgp. 106 Cancer cells that are farther 

away from the SLN could possibly develop resistance to the future chemotherapeutic treatment. 

Another feasible approach to encapsulate anthracycline salts is to prepare PLN instead. With 

the inclusion of dextran sulfate to provide counterions, PLN of doxorubicin hydrochloride were 

successfully prepared. 55 The encapsulation efficiency of doxorubicin in the PLN, depending on the 

drug payload, was generally over 70% in the presence of dextran sulfate versus approximately 40% 

in its absence. More recently, another doxorubicin-loaded PLN formulation that used a more 

lipophilic HPESO polymer (hydrolyzed and polymerized epoxidized soybean oil) was formulated. 

10 Payload over 6% was achieved. The drug releases from both PLN formulations were 

fast compared to ion pair-based SLN (see Figure 36.10). The lower drug encapsulation efficiency 

and faster drug release compared to the ion pair-based SLN indicate that unlike ion pair formation, 

the drug–polymer complexation probably does not completely neutralize all of the charges on the 

polymer molecules. These residual charges might help draw the water molecules into the lipid 

matrix to accelerate its disintegration, and they may lead to faster and more complete drug release 

from the nanoparticles. If this will lead to more effective cancer therapy still requires 

further investigations. 

36.5.4.3 Water-Soluble, Non-Ionic Drug Molecule 

36.5.4.3.1 General Principle 


All of the strategies mentioned in Section 36.4.1 rely on neutralization of the charge on the drug 

molecules. For small, non-ionic hydrophilic molecules, these strategies cannot be applied. In fact, 

researchers have already encountered similar technical challenges in the formulation of SLN for the 

treatment of non-cancer diseases. For instance, 3 0 azido-3 0 deoxythymidine (AZT) is a non-ionic 

antiviral drug with low molecular weight (MWZ267.24). This drug is too water soluble to be 

formulated into SLN. One way to overcome this difficulty is to prepare a lipophilic drug derivative. 

In the case of AZT, researchers conjugated a palmitate group to the AZT molecule, and the drug 

loading was shown as significantly improved subsequent to the increased lipophilicity. 107 

Hydrophilic cytotoxic compounds are relatively few. Both 5-FU and mitomycin-C fall into this 

category. Cisplatin and its analog carboplatin can also be considered as water-soluble small drugs. 

Until recently, only 5-FU has been attempted for SLN encapsulation. 



36.5.4.3.2 Fluorouracil 

The drug-derivatization strategy was adopted for SLN encapsulation of 5-FU. The water solubility 

of 5-FU is 12.2 mg/ml that makes it very difficult to be loaded in SLN. Wang et al. reduced the 

water solubility by conjugating two octanoyl groups to the 5-FU molecule, 90 resulting in 

3 0 ,5 0 -dioctanoyl-5-fluoro-2 0 -deoxyuridine (FuDR). This lipophilic drug derivative could be 

loaded into SLN with an encapsulation efficiency at over 90%. This strategy, however, requires 

tedious procedures of chemical synthesis. The chemical purity and identity of the final product have 

to be confirmed, and the toxicity and efficacy of the derivative also needs to be evaluated. All these 

steps demand substantial amounts of work. Drugs like 5-FU can be easily diluted and administered 

without the concern for drug precipitation such as in the cases of the lipophilic compounds, this 

probably provides very little motivation for the researchers to go through all these tedious 

procedures. However, one must keep in mind that SLN provide more than just drug solubilization 

and stabilization. They also allow controlled release, and they possibly improve the anti-tumor 

activities, side effect profiles, and biodistribution patterns of the loaded drugs. Successful development 

of SLN formulations for this class of anti-tumor drugs can still be rewarding. 

36.5.4.3.3 Overall Remarks on Cytotoxic Drug Delivery 

Cytotoxic anti-tumor drugs are a class of compounds well known for their heterogeneity in terms of 

molecular structures and physicochemical properties. Although a number of strategies have been 

devised to allow encapsulation of water-soluble salts of cytotoxic compounds, so far, only two 

anthracyclines have been successfully encapsulated. It is obvious that there is more to explore in 

this area. For low molecular weight and neutral water-soluble compounds, the major limitation of 

SLN—inefficient loading of hydrophilic molecules—is still a concern. Until recently, only limited 

efforts using the drug-derivatization strategy have been implemented. The various strategies that 

may be suitable for encapsulation of the different classes of cytotoxic compounds by SLN are 

summarized in Table 36.5. Some of the newer SLN-based systems have not been tested for their use 

for anti-tumor drug delivery (e.g., NLC and LDC), but based on their general properties, it is 

expected that they should be able to adequately handle the suggested drug types. It is expected 

that even more well-designed systems will be developed in the near future to allow more specific 

and efficient cytotoxic drug delivery. 

TABLE 36.5 

A Summary of the Possible SLN-Based Systems Currently Available That Can Be Used for 

the Delivery of Cytotoxic Drugs and Chemosensitizers 

Drug Physicochemical Properties Possible Systems 

Cytotoxic drugs Lipophilic, water-insoluble molecules SLN, NLC 

Water-soluble, ionic salts Ion-pair-SLN, LDC, PLN 

Water-soluble, hydrophilic small molecules SLN or NLC with lipophilic drug derivative 

Chemosensitizers Lipophilic, amphiphilic molecules in salt 

forms 

SLN, NLC similar to cytotoxic drugs 

SLN, classical SLN and microemulsion-based SLN; LD, lipid-drug conjugate lipid nanoparticles; PLN, polymer lipid hybrid 

nanoparticles. 


764 

36.5.5 DELIVERY OF CHEMOSENSITIZING AGENTS USING SLN 

36.5.5.1 Multidrug Resistance and Chemosensitizers 

As previously mentioned, MDR phenotype in cancer cells presents a significant obstacle to antitumor 

therapy at a cellular level. MDR was demonstrated when tumor cells that have been exposed 

to one cytotoxic agent develop cross resistance to a broad range of structurally and functionally 

unrelated compounds. 108 The cytotoxic drugs that are most frequently associated with MDR are 

typically hydrophobic, amphipathic products, including vinca alkaloids (vincristine, vinblastine), 

taxanes (paclitaxel, docetaxel), epipodophyllotoxins (etoposide, teniposide), anthracyclines (doxorubicin, 

daunorubicin, epirubicin), topotecan, and mitomycin C. 109,110 

MDR mediated by membrane-bound drug efflux transporters such as P-glycoprotein (Pgp) is 

probably the most studied form of drug resistance. 76,111 If MDR as a result of Pgp expression could 

be reduced, it would possibly increase the success rate of cancer chemotherapy. One approach is to 

use chemicals with Pgp modulatory activity to restore the drug sensitivity of cancer-demonstrating 

MDR. These agents, including verapamil, cyclosporin A, quinidine, tamoxifen, and several calmodulin 

antagonists, were identified in the 1980s, 110 and they were often referred as chemosensitizers 

or multi-drug resistance reversal agents. Newer and more potent analogs of the older chemosensitizers, 

including Valspodar (PSC833), Biricodar, and later, the more Pgp-specific synthetic 

compounds such as XR9576, 112 LY335979, 113 and Elacridar (GG918/GF120918) 114 were 

successively developed. 

36.5.5.2 Rationales for Delivery of Chemosensitizers Using SLN 

When tested in clinical settings, the earlier chemosensitizers often produced disappointing results, 

primarily because of their low affinities for Pgp. High doses of these agents are required to achieve 

the desired chemosensitizing effect, and this results in unacceptable toxicity 115 Many of these 

chemosensitizers are also substrates for other enzymes and transporters. This further leads to 

unpredictable pharmacokinetic interactions when used with other anti-tumor agents. Although 

the recently developed chemosensitizers are more potent, less toxic, and more Pgp-specific, pharmacokinetic 

interactions still occur 115–117 even when the newest agents such as GG918 were 

chosen. 118 

Like most Pgp substrates, chemosensitizers are usually hydrophobic molecules. Many chemosensitizers 

also carry basic, ionizable groups like doxorubicin. Figure 36.13a and Figure 36.13b 

show the molecular structures of two examples of chemosensitizers. These properties allow them to 

be incorporated into SLN at high efficiency. By delivering chemosensitizers using SLN, it is hoped 

MeO 

MeO 

CN CH3 C (CH2) 3 N 

CH(CH3 ) * 

2 

CH2CH2 OMe 

OMe 

OMe 

(a) Verapamil (b) GG-918 

FIGURE 36.13 Examples of molecular structures of two chemosensitizers. (a) Verapamil; and (b) GG918. 

* Basic amine group that is ionizable by protonation. 


O 

N 

H 


O 

N 

H 

N * 

OMe 

OMe


that the risk of pharmacokinetic drug interactions will be further minimized. Moreover, even 

without any surface-modifications, SLN are able to improve tumoral drug concentration (to be 

discussed later), so the chemosensitizing effects could potentially be improved, and the non-cellular 

drug resistance that is contributed by subtherapeutic tumoral drug concentration may also 

be alleviated. 

36.5.5.3 Current Use of SLN for Chemosensitizer Delivery 

So far, only a handful of chemosensitizers have been encapsulated by SLN. These include cyclosporin-A, 

verapamil, quinidine, and GG918. The cyclosporin-A SLN formulation was actually 

intended for immunomodulation, the official indication of cyclosporin-A, rather than for drug 

resistance reversal. 119 The discussion will, therefore, be focused on the other three compounds. 

36.5.5.3.1 Verapamil and Quinidine 

Verapamil and quinidine are both lipophilic molecules commonly available in ionic salt forms 

(verapamil hydrochloride and quinidine sulfate) that can both partition into a lipophilic phase when 

complexed to an anionic polymer. It was determined that the encapsulation efficiencies of the salts 

of verapamil and quinidine in lipid nanoparticles were improved from 19.6 to 69.3% and 29.5 to 

58.3%, respectively, when dextran sulfate was co-loaded to form PLN. 55 When verapamil and 

quinidine were simultaneously loaded, those encapsulation efficiencies slightly dropped to 50.3% 

and 55.8%, still significantly higher than without the addition of an anionic polymer. The drug 

release rates of both drugs from the dual-loaded PLN system were similar to when the drugs were 

separately loaded. The study shows that it is possible to simultaneously load and deliver two 

chemosensitizers by PLN with only marginal interferences between each other. This could be 

therapeutically valuable as it was reported that combinations of two chemosensitizers, even both 

at suboptimal concentrations, were able to achieve synergistic or additive effect to overcome MDR 

because of Pgp-over expression. 120 

PLN dual-loaded with verapamil and doxorubicin were also evaluated. 55 The drug loading 

properties and drug release kinetics of the dual-loaded system was similar to the single-drug 

systems. Nevertheless, the systemic cytotoxicity of verapamil is very high (LD50Z163 mg/g in 

mice). Considering this drug requires a concentration in the range of mM to mM to be effective for 

multidrug resistance (MDR) reversal, 121 there is too much risk to put it in clinical use even it is in 

encapsulated form. The study, however, reveals the possibility of using SLN-based system for 

multiple drug delivery that is critical in terms of cancer chemotherapy as drug combinations are 

typically used. 

36.5.5.3.2 Elacridar (GG918) 

GG918 has strong inhibitory activity on Pgp and breast cancer resistance protein, 122,123 two transporters 

that can contribute to cellular drug resistance. It is more potent and safer than 

verapamil. 124,125 The use of SLN for simultaneous delivery of GG918 and doxorubicin is still at 

experimental stage. 126 It was reported that the two drugs could be efficiently loaded in the same 

PLN formulation (co-loaded with HPESO polymer), and the resulting dual-drug system is effective 

against Pgp-over expressing cancer cell lines. It is noteworthy that the dual-drug system was also 

significantly more effective than when the two drugs were both in solution form or both in separate 

single-drug PLN systems. In initial results, 126 the anti-tumor effects of various formulations and 

their combinations were as follows: 

doxorubicin/GG918-PLN O free doxorubicin/GG918 in solution O doxorubicin-PLN C GG918 

solution z doxorubicin solution C GG918-PLN O doxorubicin-PLN C GG918-PLN (a mixture of 

two single-drug PLN) 


766 

In other words, synergistic effect in drug-resistant cancers may be achieved using this type of 

dual-drug SLN (or PLN) system. The drug combination appears to require proper spatial distributions 

to optimally function. When the two drugs are separately encapsulated, the synergistic 

effect is lost. These findings have laid the foundation for using SLN-based system for combinational 

drug chemotherapy. 

36.5.5.3.3 Overall Remarks on Chemosensitizer Delivery 

The study of SLN systems for the delivery of chemosensitizers is still in its infancy. Only limited 

information in this field has been revealed thus far. Given the good lipophilicities of chemosensitizers 

in general, it is not surprising to see that the few chemosensitizers studied until now are well 

incorporated in PLN. It is expected that other free-base chemosensitizers can be loaded into even 

unmodified, classical SLN (See Table 36.5). Formulation of PLN dual-loaded with two chemosensitizers 

or a chemosensitizer/cytotoxic agent combination was also shown to be possible. This, 

once again, demonstrates the versatility of SLN-based delivery systems. 

36.5.6 EFFECT OF SLN FORMULATION ON IN VITRO ANTI-TUMOR TOXICITY 


The primary goals of cytotoxic drug therapy are to kill as many cancer cells as possible and to 

prevent their proliferation. Therefore, it is important to ensure that the drug released from a SLN 

system preserves its cytotoxicity, and the drug delivered by SLN is at least as cytotoxic to cancer 

cells as the conventional free drug solution. These two issues appear identical but are, in fact, quite 

different. Drug releases from SLN are rarely 100% complete. In fact, in many SLN formulations to 

be discussed later, only fractions of the loaded anti-tumor drugs are releasable. However, these 

formulations all demonstrated cancer cytotoxicity comparable or even superior to the corresponding 

free drugs. These findings suggest that, in addition to simply unloading the drug 

outside the cancer cells and killing the cells with this fraction of free, released drug, the portion 

of drug that remains associated with the SLN is still cytotoxic to various extents. SLN quite 

possibly have additional mechanisms to deliver this unreleased drug to achieve cancer cell kill 

or growth suppression. It is important to evaluate the cytotoxicity of a drug-loaded SLN formulation 

as an integral unit and, whenever possible, to also investigate the underlying cytotoxicity 

mechanisms that derive from the use of SLN. 

Only until recently, a limited number of in vitro cytotoxicity studies on anti-tumor drug-loaded 

SLN have been conducted. Miglietta et al. performed trypan blue exclusion assays on breast cancer 

cell line MCF-7 and leukemia cell lines (HL60) treated with SLN incorporating doxorubicin or 

paclitaxel for 72 h. 83 The results of the assays demonstrated improved cytotoxicity with both 

formulations compared to the corresponding free drugs. A similar study using the same methodology 

was later carried out to evaluate the cytotoxicity of SLN formulations carrying doxorubicin, 

paclitaxel, or cholesteryl butyrate on a colorectal cancer cell line HT-28. 62 SLN of doxorubicin and 

cholesteryl butyrate were both significantly more cytotoxic than the corresponding free drugs as 

indicated by the lower drug concentrations needed for the SLN to achieve 50% cell non-viability 

(IC 50 of SLN versus solution: doxorubicin: 81.87 nM versus 126.57 nM, butyrate: 0.3 mM versus 

O 0.6 mM;). Paclitaxel SLN were similarly cytotoxic compared to the drug solution. This was 

possibly caused by the extremely slow release of paclitaxel from the SLN. 82 

One interesting finding in Serpe et al.’s study is about the combination use of anti-tumor 

drugs. 62 The authors reported that the combination of low concentrations of cholesteryl butyrateloaded 

SLN and doxorubicin or paclitaxel in solution exerted synergistic reduction in cancer cell 

viability, whereas no such reduction was demonstrated using the same combinations of drugs both 

as solution. Although the cytotoxicities appear low and only trypan blue exclusion assays were 

employed, the findings nonetheless support the possibility that combination drug therapy involving 

SLN formulations could be more effective than the conventional free drug treatment. 



Some nanoparticle systems and liposomal formulations have been shown to be effective against 

cancer cells with MDR phenotype. 127–130 These systems were all made of lipids or lipophilic 

substances, and many of them were coated with non-ionic surfactants. Also with lipid-based 

formulations that often use non-ionic surfactants, it is reasonable to expect MDR reversal activities 

in at least some SLN systems. In the aforementioned studies, even though HL-60 and HT-28 cell 

lines are not specifically drug-resistant cell lines, they are moderately refractory to doxorubicin, and 

the SLN formulations were still able to effectively treat them. In another study, PLN dispersed 

using Pluronic F68, a non-ionic block co-polymer, were able to enhance the cytotoxicity of doxorubicin 

in Pgp-over expressing breast cancer cell line MDA435/LCC6/MDR1. 10 Doxorubicinloaded 

PLN also led to significantly stronger reductions in cell viability in trypan blue exclusion 

assays than free doxorubicin. In addition, the use of clonogenic assays, a more sensitive method to 

evaluate the effectiveness of cancer proliferation suppression, showed a larger than 8-fold increase 

in the anti-tumor activity using PLN formulation compared to free doxorubicin (Figure 36.14). It is 

interesting to observe that blank particles did not enhance the anti-tumor activity of doxorubicin 

solution, and doxorubicin–polymer (Dox–HPESO) was not more effective than free doxorubicin. 

The findings of this study further reveal the potential of SLN for drug-resistant cancer treatment, 

and they support the earlier claim that encapsulated cytotoxic compound is still effective. In fact, a 

cytotoxic compound may need to be encapsulated to achieve that enhanced activity in drugresistant 

cancer. 

Initial work has been undertaken to delineate the underlying MDR reversal mechanisms by 

SLN. 11 Increased drug uptake and more noticeably, prolonged drug retention, were observed when 

doxorubicin-loaded PLN were used instead of free drug. This was further supported by fluorescence 

microscope images. 11 In Figure 36.15, cancer cells were treated for 2 h and re-incubated in fresh 

medium for an additional 2 h. A high level of doxorubicin that has fluorescent property was 

intracellularly detected when the drug was delivered in form of PLN. Another fluorescence 

microscopy study also revealed that lipids were associated with the loaded drug molecules in the 

cells. All these findings suggest that some of the nanoparticles may carry doxorubicin into the cell, 

Normalized PE 

1 

0.1 

0.01 

0.001 

Dox 

Dox-HPESO 

Dox-PLN 

Dox + Blank PLN 

0 2 4 6 8 10 

Dox concentration (μg/ml) 

FIGURE 36.14 Clonogenic assay experiments for the toxicity of Dox human P-glycoprotein overexpressing 

MDA435/LCC6/MDR1 breast cancer cell line. The normalized plating efficiencies (normalized PE) after 4 h 

exposure to doxorubicin (Dox) solution, Dox-polymer (Dox-HPESO) aggregates, Dox-loaded lipid nanoparticles 

(Dox-PLN), or Dox solution C blank PLN are shown. Results are expressed as mean G SD of the 

measurements obtained in three separate experiments (nZ6 in each experiment). *, p ! 0.05, significantly 

different from Dox solution group. (Adapted from Wong, H. L. et al., Pharm. Res., in press.) 


768 

FIGURE 36.15 Fluoroescence microscope images demonstrating the effect of a PLN formulation containing 

doxorubicin on drug retention and intracellular drug distribution in Pgp-over expressing breast cancer cells 

MDA435/LCC6/MDR1. Cells were incubated with nanoparticles containing doxorubicin for 2 h, washed, and 

re-incubated in fresh medium for two additional hours. l ex Z540 nm and l em Z590 nm. Magnification of 

objective Z100!. Bar represents 20 mm. (Adapted from Wong, H. L. et al., Pharmacol. Exp. Ther., in press.) 

possibly by endocytosis, bypassing the membrane-bound drug efflux mechanisms. These drugloaded 

nanoparticles could be staying intracellularly to serve as mini-drug depot systems and 

chronically inflicting damages to cancer cells. 

36.5.6.1 In Vivo Efficacy of SLN Delivered Antineoplastic Agents 

In terms of cancer therapy, particularly with regard to solid tumors, in vitro data frequently may 

only serve as a good starting point. It is because there are several drug resistance mechanisms, 

sometimes referred as non-cellular mechanisms, that exist only in in vivo situations. A treatment 

with strong in vitro anti-tumor activity thus maynotbeworkingintumor-bearinganimals. 

However, information on the in vivo efficacy of SLN formulations of anti-tumor drugs remains 

limited. Most studies that included animal models focused on drug biodistribution. To our knowledge, 

only one study with in vivo efficacy component has been published. In this study, mice model 

xenografted with HT-28 tumor was used. 62 Tumors treated with the SLN formulation of SN-38 took 

longer or comparable time to reach the cut off tumor weight at lower drug dose. The study provided 

the first evidence that the ability of SLN to enhance anti-tumor drug effectiveness is not limited in 

cultured cells, but may also be applicable to in vivo systems. 

36.5.6.2 Improved Pharmacokinetics and Drug Distribution 


As discussed in Section 36.2.3, SLN can be coated with stealth agents such as PEG to minimize the 

clearance by RES to achieve extended systemic circulation time. Using stearic acid PEG2000, longcirculating 

SLN formulations of doxorubicin and paclitaxel have been formulated. 9,84 A number of 

pharmacokinetic and biodistribution studies were based on these SLN systems. 

In principle, the unmodified, non-stealth SLN should be rapidly cleared from the systemic 

circulation by RES. In reality, unmodified SLN were able to remain in the bloodstream for long 

times in the animal studies performed. Table 36.6 summarizes the area-under-the-curve (AUC) 



TABLE 36.6 

A Comparison of the Area-Under-the-Curve Values of Anti-Cancer Drugs Administered in 

Solution to Drugs Encapsulated in Unmodified SLN or Long-Circulating SLN 

Study Group a 

Drug b 

AUC c 

Funaro Yang Wang Zara 99 Zara 02 Zara 02 

DOX CTT FuDR DOX DOX IDA 

Drug solution 83.7 66 44.5 98.0 47.6 31.4, 108.5 d 

Unmodified SLN 814.5 324 138.4 1713.4 123.1 674.2, 458.9 

Long-circulating SLN 1121.1 — — — 259.7 e 

318.2 

433.3 

— 

a Groups named under the first authors of the studied. 

b DOX, doxorubicin; CTT, camptothecin; FuDR, 3 0 ,5 0 -dioctanoyl-5-fluoro-2 0 -deoxyuridine; IDA, idarubicin. 

c AUC, area under curve, units in min.(g.ml K1 except the studies by Yang (h.ng.g K1 ) and Wang (h.mg.ml K1 ). 

d First figure and second figure correspond to AUC of IDA after intravenous and duodenal administration, respectively. 

e SLN with three different levels of PEG2000-stearic acid.First, second, and third figures correspond to 0.15%, 0.30%, and 

0.45% PEG2000 content in lipid microemulsion for SLN preparation. 

values of several anti-tumor compounds obtained in pharmacokinetic studies. Drugs in solution, 

unmodified SLN, or long-circulating SLN were administered. In all of the studies listed, unmodified 

SLN were able to increase the AUC values by 3-fold up to over 20-fold, indicating low RES 

clearance. Nevertheless, long-circulating SLN still have an advantage over the unmodified SLN. 

In the two studies that compared unmodified and long-circulating formulations, the AUC of drugs 

delivered by the long-circulating SLN were approximately 25% to over 3-fold higher. 

In addition to surface modification, the routes of administration also appear to affect the 

pharmacokinetics and drug biodistribution. In the etoposide-SLN study, 80 improvements in 

tumoral drug accumulation were observed when the etoposide formulations were intraperitoneally 

or subcutaneously injected. Subcutaneous administration even led to multiple-fold increases in the 

tumor drug concentrations 24 h post-injection. The authors suggested that the slower and progressive 

penetration of the nanoparticles (and the loaded drug) from the subcutaneous injection 

site into the tumor may result in more favorable patterns of drug distribution. In the study by Zara 

et al., 88 duodenal administration of idarubicin-loaded SLN also led to a higher AUC than when the 

SLN were intravenously administered. It is evident that the biodistribution of an anti-tumor drug 

delivered by SLN may be further manipulated to achieve the desired therapeutic goal. The manner 

by which a SLN drug formulation is administered will be a key aspect to consider when designing 

animal or clinical studies of SLN for anti-tumor drug delivery. 

Regardless of how long an anti-tumor drug can stay in the systemic circulation, the drug has to 

finally accumulate in the tumor and avoid the non-cancerous tissues to exercise its therapeutic 

effects without causing toxicity to the patient. To confirm the tumor drug accumulation, an in vivo 

tumor model is required. Only one SLN study employing tumor-bearing animals has been 

performed. 80 Nearly 67% and 30% increases of tumor drug concentrations were measured 1 h 

and 24 h post-injection, respectively. This is promising in terms of cancer chemotherapy. More 

SLN biodistribution studies involving tumor models will further strengthen the case. 

In many of the biodistribution studies of anti-tumor drug-loaded SLN, even without surface 

modifications, the drug accumulations in the liver, spleen, and kidney, three RES organs, were 

reduced. 8,9,80,88 The role of RES clearance in the biodistribution of SLN may need to be 

re-examined. It is also noteworthy to point out the high brain concentrations of SLN delivered 

drugs in many of these studies. Cytotoxic drugs do not usually accumulate in the brain tissue 


770 

because of the presence of the blood–brain barrier. This barrier presents a serious challenge to 

chemotherapeutic treatment of central nervous system tumors (refer to Chapter 17). The most 

prominent mechanism leading to the blood–brain barrier is the high expression of Pgp. 

The ability of SLN to carry drugs across the Pgp-rich blood–brain barrier is consistent with the 

previously described findings of SLN in Pgp-over expressing cancer cell lines (Section 36.5.3). This 

Pgp-bypassing feature of SLN may be useful for cancer chemotherapy if properly exploited. 

In most of the SLN biodistribution studies, the drug concentrations in the heart were significantly 

lower using SLN formulations. 8,9,80,88 This is important for anthracycline delivery as this 

class of anti-tumor compounds are notorious in causing irreversible cardiotoxicity. By selectively 

reducing the anthracycline concentration in the cardiac tissues, higher dose intensity may be given 

to achieve better therapeutic outcomes. 131 

Overall, SLN are useful for the improvement of the pharmacokinetics and biodistribution 

profiles of anti-tumor drugs. Drugs administered tend to stay longer, penetrate the tumor sites 

better, and avoid some vulnerable organs. These may be further improved by the use of longcirculating 

SLN systems. 



SLN are a promising drug delivery device for rational chemotherapy. After just slightly longer than 

a decade of time, SLN already have gone through several stages of development. Although the 

ingredients of SLN are relatively simple—basically only lipids and surfactants—there are a diverse 

number of parameters involved in the preparation processes that can be manipulated to alter the 

SLN properties. These include the physicochemical characteristics of SLN compositions, production 

methods, and storage conditions. Significant progresses in the improvement of drug 

loading capacity, mass-production, modulation of release kinetics, stability in storage, bioavailability, 

and active targeting have been made. In addition, the early, classical SLN that were made of 

highly crystallized lipids have quickly evolved into several variations, including microemulsionbased 

SLN, NLC, LDC, PLN, and surface-modified SLN. Each of these variations corrects a 

specified set of problems commonly encountered during the early stage of SLN development. 

Currently not only limited to lipophilic compounds, hydrophilic and charged agents can also be 

efficiently encapsulated into SLN-based systems, stored with good stability, and released under a 

certain degree of control. This makes SLN an excellent choice for the delivery of anti-tumor drugs, 

a highly heterogeneous class of drugs with diverse molecular structures and physical properties and 

are usually chemical reactive and strongly toxic. 

This relatively new class of delivery systems still has a lot of untapped potential as well as 

limitations to overcome. Because of the complexity of such drug delivery systems and complex 

interactions existed among various causal factors, the fundamental problems hampering the practical 

application of SLN such as drug loading and release mechanisms, the correlation between the 

constituents of SLN and the properties of SLN have not been completely solved yet. The current 

development of SLN still mainly depends on the trial-and-error approach that leads to occasional 

controversial results. It is expected that with better understanding of the structure of SLN and their 

mechanisms of drug loading and release, the overall performance of SLN can be further optimized. 

The biological aspects of SLN for cancer treatment, including SLN-cell interactions and in vivo 

SLN-tumor interactions are areas largely unexplored. In addition, there are a number of new 

applications of SLN that have just started developing, e.g., the use of SLN for delivery of drug 

combinations such as cytotoxic drug—chemosensitizer. The potentials of SLN for gene therapy and 

molecular targeting are also emerging. 84,132 It will not be surprising to witness significant research 

progress in the field of SLN in the near future, providing more efficacious, less toxic, and more 

specific chemotherapy treatments for cancer. 



REFERENCES 

1. Müller, R. H., Colloidal carriers for controlled drug delivery and targeting: Modification, characterization 

and in vivo distribution, CRC Press, 45–55, 1991. 

2. Wissing, S. A., Kayser, O., and Müller, R. H., Solid lipid nanoparticles for parenteral drug delivery, 


3. Müller, R. H. et al., Solid lipid nanoparticles (SLN)-An alternative colloidal carrier system for 

controlled drug delivery, Eur. Pharm. Biopharm., 41, 62–69, 1995. 

4. Schwarz, C. et al., Solid lipid nanoparticles (SLN) for controlled drug delivery. I. Production, 

characterization and sterilization, J. Control. Release, 30, 83–96, 1994. 

5. Almeida, A. J., Runge, S., and Müller, R. H., Peptide-loaded solid lipid nanoparticles (SLN): 

Influence of production parameters, Int. J. Pharm., 149, 255–265, 1997. 

6. Müller, R. H., Redtke, M., and Wissing, S. A., Solid lipid nanoparticles (SLN) and nanostructured 

lipid carriers (NLC) in cosmetic and dermatological preparations, Adv. Drug Deliv. Rev., 54, 

S131–S155, 2002. 

7. Wissing, S. A. and Müller, R. H., Solid lipid nanoparticles as carrier for sunscreens: In vitro release 

and in vivo skin penetration, J. Control. Release, 81, 345–355, 2002. 

8. Zara, G. P. et al., Intravenous administration to rabbits of non-stealth and stealth doxorubicin-loaded 

solid lipid nanoparticles at increasing concentration of stealth agent: Pharmacokinetics and distribution 

in brain and other tissues, J. Drug Target., 10, 327–335, 2002. 

9. Fundarò, A. et al., Non-stealth and stealth solid lipid nanoparticles (SLN) carrying doxorubicin: 

Pharmacokinetics and tissue distribution after I.V. administration to rat, Pharmacol. Res., 42, 

337–343, 2000. 

10. Wong, H. L. et al., A new polymer–lipid hydrid nanoparticle system increases cytotoxicity of doxorubicin 

against multidrug resistant human breast cancer cells, Pharm. Res., 23, 7, 1574–1585, 2006. 

11. Wong, H. et al., A Mechanistic study of enhanced doxorubicin uptake and retention in multidrug 

resistant breast cancer cells using a polymer–lipid hybrid nanoparticle (PLN) system, J. Pharmcol. 

Exp. Ther., (in press). 

12. Westesen, K., Siekmann, B., and Koch, M. H. J., Investigations on the physical state of lipid 

nanoparticles by synchrotron radiation X-ray diffraction, Int. J. Pharm., 93, 189–199, 1993. 

13. Dingler, A. and Gohla, S., Production of solid lipoid nanoparticles (SLN): Scaling up feasibilities, 

J. Microencapsul., 19, 11–16, 2002. 

14. Marengo, E. et al., Scale-up of the preparation process of solid lipid nanoparticles. Part I, Int. 

J. Pharm., 205, 3–13, 2000. 

15. Jenning, V., Lippacher, A., and Gohla, S. H., Medium scale production of solid lipid nanoparticles 

(SLN) by high pressure homogenization, J. Microencapsul., 19, 1–10, 2002. 

16. Gohla, S. H. and Dingler, A., Scaling up feasibility of the production of solid lipid nanoparticles 

(SLN), Pharmazie, 56, 61–63, 2001. 


1818–1822, 2004. 

18. Moses, M. A., Brem, H., and Langer, R., Advancing the field of drug delivery: Taking aim at cancer, 

Cancer Cell, 4, 337–341, 2003. 

19. Jenning, V. and Gohla, S., Comparison of wax and glyceride solid lipid nanoparticles (SLN), Int. 

J. Pharm., 196, 219–222, 2000. 

20. Jenning, V. and Gohla, S. H., Encapsulation of retinoids in solid lipid nanoparticles (SLN), 

J. Microencapsul., 18, 149–158, 2001. 

21. Bunjes, H. et al., Incorporation of the model drug ubidecarenone into solid lipid nanoparticles, 

Pharm. Res., 18, 287–293, 2001. 

22. Aulton, M. E., Pharamaceutics: The Science of Dosage form Design., Churchill Livingstone, Edinburgh. 

81–118, 1988. 

23. Bunjes, H., Koch, M. H. J., and Westesen, K., Influence of emulsifiers on the crystallization of solid 

lipid nanoparticles, J. Pharm. Sci., 92, 1509–1520, 2002. 

24. Aquilano, D., Cavalli, R., and Gasco, M. R., Solid lipospheres obtained from hot microemulsions in 

the presence of different concentration of cosurfactant: The crystallinizatiion of stearic acid polymorphs, 

Thermochim. Acta., 230, 29–37, 1993. 


772 


25. Olbrich, C. and Müller, R. H., Enzymatic degradation of SLN-effect of surfactant and surfactant 

mixtures, Int. J. Pharm., 180, 31–39, 1999. 

26. Olbrich, C., Kayser, O., and Müller, R. H., Enzymatic degradation of Dynasan 114 SLN—effect of 

surfactants and particle size, J. Nanoparticle Res., 4, 121–129, 2002. 



SMANCS, Cancer Res., 6, 193–210, 1986. 

28. Müller, R. H., Mader, K., and Gohla, S., Solid lipid nanoparticles (SLN) for controlled drug 

delivery—a review of the state of art, Eur. J. Pharm. Biopharm., 50, 161–177, 2000. 

29. Kawashima, Y., Nanoparticulate systems for improved drug delivery, Adv. Drug Deliv. Rev., 47, 

1–2, 2001. 

30. Mehnert, W. and Mader, K., Solid lipid nanoparticles, production, characterization and applications, 


31. Westesen, K., Bunjes, H., and Koch, M. H. J., Physicochemical characterization of lipid nanoparticles 

and evaluation of their drug loading capacity and sustained release potential, J. Control. 

Release, 48, 223–236, 1997. 

32. Shahgaldian, P. et al., Para-acyl-calix-arene based solid lipid nanoparticles (SLNs): A detailed study 

of preparation and stability parameters, Int. J. Pharm., 253, 23–28, 2003. 

33. Bummer, P. M., Physical chemical considerations of lipid-based oral drug delivery-solid lipid 

nanoparticles, Crit. Rev. Ther. Drug Carrier Syst., 21, 1–20, 2004. 

34. Manjunath, K., Reddy, J. S., and Venkateswarlu, V., Solid lipid nanoparticles as drug delivery 

systems, Methods Find. Exp. Clin. Pharmacol., 27, 127–144, 2005. 

35. Li, Y., Taulier, N., Rauth, A. M., and Wu, X. Y., American Association of Pharmaceutical Scientists 

(AAPS) conference. TN, USA, Nov. 7–11, 2005. 

36. Chapman, D., The polymorphism of glycerides, Chem. Rev., 62, 433–456, 1961. 

37. Bunjes, H., Koch, M. H. J., and Westesen, K., Influence of emulsifiers on the crystallization of solid 

lipid nanoparticles, J. Pharm. Sci., 92, 1509–1520, 2003. 

38. Hou, D. et al., The production and characteristics of solid lipid nanoparticles (SLNs), Biomaterials, 

24, 1781–1785, 2003. 

39. Venkateswarlu, V. and Manjunath, K., Preparation, characterization and in vitro release kinetics of 

clozapine solid lipid nanoparticles, J. Control. Release, 95, 627–638, 2004. 

40. Olbrich, C., Kayser, O., and Müller, R. H., Lipase degradation of Dynasan 114 and 116 solid lipid 

nanoparticles (SLN)-effect of surfactants, storage time and crystallinity, Int. J. Pharm., 237, 

119–128, 2002. 

41. Souto, E. B. et al., Development of a controlled release formulation based on SLN and NLC for 

topical clotrimazole delivery, Int. J. Pharm., 278, 71–77, 2004. 

42. Müller, R. H., Radtke, M., and Wissing, S. A., Nanostructured lipid matrices for improved microencapsulation 

of drugs, Int. J. Pharm., 242, 121–128, 2002. 

43. Jenning, V., Thunemann, A. F., and Gohla, S. H., Characterization of a novel solid lipid nanoparticle 

carrier system based on binary mixtures of liquid and solid lipids, Int. J. Pharm., 199, 167–177, 

2000. 

44. Jenning, V., Korting, M. S., and Gohla, S., Vitamin A-loaded solid lipid nanoparticles for topical use: 

Drug release properties, J. Control. Release, 66, 115–126, 2000. 

45. Tabatt, K. et al., Effect of cationic lipid and matrix lipid composition on solid lipid nanoparticlemediated 

gene transfer, Eur. J. Pharm. Biopharm., 57, 155–162, 2004. 

46. Chen, D. N. et al., In vitro and in vivo study of two types of ion-circulating solid lipid nanoparticles 

containing paclitaxel, Chem. Pharm. Bull., 49, 1444–1447, 2001. 

47. Bocca, C. et al., Phagocytic uptake of fluorescent stealth and non-stealth solid lipid nanoparticles, Int. 

J. Pharm., 175, 185–193, 1998. 

48. Gref, R. et al., The controlled intravenous delivery of drugs using PEG-coated sterically stabilized 

nanoparticles, Adv. Drug Deliv. Rev., 16, 215–233, 1995. 

49. Illum, L. et al., The organ distribution and circulation time of intravenous injected colloidal carriers 

sterically stabilized with a block copolymer–Poloxamine 908, Life Sci., 40, 367–374, 1987. 



50. Jores, K., Mehnert, W., and Mader, K., Physicochemical investigations on solid lipid nanoparticles 

and on oil-loaded solid lipid nanoparticles: A nuclear magnetic resonance and electron spin resonance 

study, Pharm. Res., 20, 1274–1283, 2003. 

51. Jores, K. et al., Investigation on the structure of solid lipid nanoparticles (SLN) and oil-loaded solid 

lipid nanoparticles by photon correlation spectroscopy, field-flow fractionation and transmission 

electron microscopy, J. Control. Release, 95, 217–227, 2004. 

52. zur Mühlen, A., Schwarz, C., and Mehnert, W., Solid lipid nanoparticles for controlled drug 

delivery—drug release and release mechanism, Eur. J. Pharm. Biopharm., 45, 149–155, 1998. 

53. zur Mühlen, A. et al., Atomic force microscopy studies of solid lipid nanoparticles, Pharm. Res., 13, 

1411–1416, 1996. 

54. Jenning, V., Mader, K., and Gohla, S. H., Solid lipid nanoparticles (SLNTM) based on binary 

mixtures of liquid and solid lipids: A 1H-NMR study, Int. J. Pharm., 205, 15–21, 2000. 

55. Wong, H. et al., Development of solid lipid nanoparticles containing ionically complexed chemometherapeutic 

drugs and chemosensitizers, Int. J. Pharm., 93, 1993–2008, 2004. 

56. Li, Y., Taulier, N., Rauth, A. M., and Wu, X. Y., Screening of lipid carriers and characterization of 

drug–polymer–lipid interactions for the rational design of polymer–lipid hybrid nanoparticles 

(PLN), Pharm. Res., (Submitted). 

57. Müller, R. H. et al., Solid lipid nanoparticles (SLN)-an alternative colloidal carrier system for 

controlled drug delivery, Eur. J. Pharm. Biopharm., 41, 62–69, 1995. 

58. Gasco, M. R., Method for producing solid lipid microspheres having a narrow size distribution, US 

Patent No. 5250236, 1993. 

59. Trotta, M., Debernardi, F., and Caputo, O., Preparation of solid lipid nanoparticles by a solvent 

emulsification-diffusion technique, Int. J. Pharm., 257, 153–160, 2003. 

60. Hu, F. Q. et al., Preparation of solid lipid nanoparticles with clobetasol propionate by a novel solvent 

diffusion method in aqueous systems and physicochemical characterization, Int. J. Pharm., 239, 

121–128, 2002. 

61. Cortesi, R. et al., Production of lipospheres as carriers for bioactive compounds, Biomaterials, 23, 

2283–2294, 2002. 

62. Serpe, L. et al., Cytotoxicity of anticancer drugs incorporated in solid lipid nanoparticles on HT-29 

colorectal cancer cell line, Eur. J. Pharm. Biopharm., 58, 673–680, 2004. 

63. Cavalli, R. et al., Solid lipid nanoparticles (SLN) as ocular delivery system for tobramycin, Int. 

J. Pharm., 238, 241–245, 2002. 

64. Cavalli, R. et al., Preparation and evaluation in vitro of colloidal lipospheres containing pilocarpine 

as ion pair, Int. J. Pharm., 117, 243–246, 1995. 

65. Gasco, M. R., Cavalli, R., and Carlotti, M. E., Timolol in lipospheres, Pharmazie, 47, 119–121, 

1992. 

66. Cavalli, R. et al., Sterilization and freeze-drying of drug-free and drug loaded solid lipid nanoparticles, 

Int. J. Pharm., 148, 47–54, 1997. 

67. Shahgaldian, P. et al., A study of the freeze-drying conditions of calixarene based solid lipid 

nanoparticles, Eur. J. Pharm. Biopharm., 55, 181–184, 2003. 

68. Zimmermann, E., Müller, R. H., and Mäder, K., Influence of different parameters on reconstitution of 

lyophilized SLN, Int. J. Pharm., 196, 211–213, 2000. 

69. Schwarz, C. and Mehnert, W., Freeze-drying of drug-free and drug-loaded solid lipid nanoparticles 

(SLN), Int. J. Pharm., 157, 171–179, 1997. 

70. Freitas, C. and Müller, R. H., Effect of light and temperature on zeta potential and physical stability 

in solid lipid nanoparticle (SLNTM) dispersions, Int. J. Pharm., 168, 221–229, 1998. 

71. Martin, A., Physical Pharmacy: Physical Chemical Principles in the Pharmaceutical Sciences, 4th 

ed., Lea & Febiger, Philadelphia, PA pp. 362–388, 1993. 

72. Freitas, C. and Müller, R. H., Stability determination of solid lipid nanoparticles (SLN) in aqueous 

dispersion after addition of electrolyte, J. Microencapsul., 16, 59–71, 1999. 

73. Skeel, R. T., Selection of treatment for the patient with cancer, In Handbook of Cancer 

Chemotherapy, Skeel, R. T., Ed., Lippincott, Williams, and Wilkins, Philadelphia, PA, pp. 46–52, 

2003. 

74. Tipton, J. M., Side effects of cancer chemotherapy, In Handbook of Cancer Chemotherapy, Skeel, 

R. T., Ed., Lippincott, Williams, and Wilkins, Philadelphia, PA, pp. 561–580, 2003. 


774 


75. Gottesman, M. M., Mechanisms of cancer drug resistance, Ann. Rev. Med., 53, 615–627, 2002. 

76. Endicott, J. A. and Ling, V., The biochemistry of P-glycoprotein-mediated multidrug resistance, 

Ann. Rev. Biochem., 58, 137–171, 1989. 

77. Lim, S. J. and Kim, C. K., Formulation parameters determining the physicochemical characteristics 

of solid lipid nanoparticles loaded with all-trans retinoic acid, Int. J. Pharm., 243, 135–146, 2002. 

78. Yang, S. C. et al., Body distribution in mice of intravenously injected camptothecin solid lipid 

nanoparticles and targeting effect on brain, J. Control. Release, 59, 299–307, 1999. 

79. Yang, S. C. and Zhu, J. B., Preparation and characterization of camptothecan solid lipid nanoparticles, 

Drug. Dev. Ind. Pharm., 28, 265–274, 2002. 

80. Reddy, L. H. et al., Influence of administration route on tumor uptake and biodistribution of etoposide 

loaded solid nanoparticles in Dalton’s lymphoma tumor bearing mice, J. Control. Release, 105, 

185–198, 2005. 

81. Williams, J. et al., Nanoparticle drug delivery system for intravenous delivery of topoisomerase 

inhibitors, J. Control. Release, 91, 167–172, 2003. 

82. Cavalli, R., Caputo, O., and Gasco, M. R., Preparation and characterization of solid lipid nanospheres 

containing paclitaxel, Eur. J. Pharm. Sci., 10, 305–309, 2000. 

83. Miglietta, A. et al., Cellular uptake and cytotoxicity of solid lipid nanospheres (SLN) incorporating 

doxorubicin or paclitaxel, Int. J. Pharmaceut., 210, 61–67, 2000. 

84. Stevens, P. J., Sekido, M., and Lee, R. J., A folate-receptor-targeted lipid nanoparticle formulation 

for a lipophilic paclitaxel prodrug, Pharm. Res., 21, 2153–2157, 2004. 

85. Videira, M. A., Almeida, A. J., and Muller, R. H., Formulation and physical stability assessment of 

SLN containing paclitaxel. Incorporation of paclitaxel in SLN: Assessment of drug–lipid interaction, 

Proc. 3rd World Meeting Pharm. Technol., 453–454, 2000. 

86. Chen, D. B. et al., In vitro and in vivo study of two types of long-circulating solid lipid nanoparticles 

containing paclitaxel, Chem. Pharm. Bull., 49, 1444–1447, 2001. 

87. Cavalli, R., Caputo, O., and Gasco, M. R., Solid lipospheres of doxorubicin and idarubicin, Int. 

J. Pharm., 89, R9–R12, 1993. 

88. Zara, G. P. et al., Pharmacokinetics and tissue distribution of idarubicin-loaded solid lipid nanoparticles 

after duodenal administration to rats, J. Pharm. Sci., 91, 1324–1333, 2002. 

89. Zara, G. P. et al., Pharmacokinetics of doxorubicin incorporated in solid lipid nanospheres (SLN), 

Pharm. Res., 40, 281–286, 1989. 

90. Wang, J. X., Sun, X., and Zhang, Z. R., Enhanced brain targeting by synthesis of 3-,5-dioctanoyl-5fluoro-2’-deoxyuridine 

and incorporation into solid lipid nanoparticles, Eur. J. Pharm. Biopharm., 

54, 285–290, 2002. 

91. Pratt, W. B. et al., The Anticancer Drugs, Oxford University Press, New York. 

92. Wall, M. E. et al., Plant antitumor agents. I. The isolation and structure of camptothecin, a novel 

alkaloidal leukemia and tumour inhibitor from camptotheca acuminate, J. Am. Chem. Soc., 88, 

3888–3890, 1966. 

93. Gottlieb, J. A. et al., Preliminary pharmacologic and clinical evaluatin of camptothecin sodium 

(NSC-100880), Cancer Chemother. Rep., 54, 461–470, 1970. 

94. Takimoto, C. H. and Arbuck, S. G., Topoisomerase I targeting agents: The camptothecins, In Cancer 

Chemotherapy and Biotherapy: Principles and Practice, Chabner, B. A. and Longo, D. L., Eds., 

Lippincott, Williams, and Wilkins, Philadelphia, PA, pp. 579–646, 2001. 

95. Rowinsky, E. K. and Donehower, R. C., Paclitaxel (taxol), N. Engl. J. Med., 332, 1004–1014, 1995. 

96. Singla, A. K., Garg, A., and Aggarwal, D., Paclitaxel and its formulations, Int. J. Pharm., 235, 

179–192, 2002. 

97. Constantinides, P. P. et al., Formulation development and antitumor activity of a filter-sterilizable 

emulsion of paclitaxel, Pharm. Res., 17, 175–182, 2000. 

98. Kan, P. et al., Development of nonionic surfactant/phospholipid o/w emulsion as a paclitaxel 

delivery system, J. Controll. Release, 58, 271–278, 1999. 

99. Bernsdorff, C., Reszka, R., and Winter, R., Interaction of the anticancer agent taxol (paclitaxel) with 

phopholipid layers, J. Biomed. Mater. Res., 46, 141–149, 1999. 

100. Clarke, S. J. and Rivory, L. P., Clinical pharmacokinetics of docetaxel, Clin. Pharmacokinet., 36, 

99–114, 1999. 



101. Ojima, I. and Geney, R., BMS-184476 Bristol-Myers Squibb, Curr. Opin. Investig. Drugs, 4, 

732–736, 2003. 

102. Baldwin, E. L. and Osheroff, N., Etoposide, topoisomerase II and cancer, Curr, Med. Chem. Anti- 

Cancer Agents, 5, 363–372, 2005. 

103. Brisaert, M., Gabriels, M., and Plaizier-Vercammen, J., Investigation of the chemical stability of an 

erythromycin-tretionin lotion by the use of an optimization system, Int. J. Pharm., 243, 135–146, 2000. 

104. Pouillart, P. R., Role of butyric acid and its derivatives in the treatment of colorectal cancer and 

hemoglobinopathies, Life Sci., 63, 1739–1760, 1998. 

105. Pellizzaro, C. et al., Cholesteryl butyrate in solid lipid nanoparticles as an alternative approach for 

butyric acid delivery, Anticancer Res., 19 (5B), 3921–3926, 1999. 

106. Hahn, S. M. et al., A multidrug-resistant breast cancer line induced by weekly exposure to doxorubicin, 

Int. J. Oncol., 14, 273–279, 1999. 

107. Heiati, H. et al., Solid lipid nanoparticles as drug carriers. I. Incorporation and retention of the 

lipphilic prodrug 3’-azido-3’-deoxythymidine palmitate, Int. J. Pharm., 146, 123–131, 1997. 

108. Biedler, J. L. and Riehm, H., Cellular resistance to actinomycin D in Chinese hamster cells in vitro: 

Cross-resistance, radioautographic, and cytogenic studies, Cancer Res., 30, 1174–1184, 1970. 

109. Ambudkar, S. V., Dey, S., and Hrycyna, C. A., Biochemical, cellular and pharmacological aspects of 

the multidrug transporter, Ann. Rev. Pharmacol. Toxicol., 39, 361–398, 1999. 

110. Krishna, R. and Mayer, L. D., Multidrug resistance (MDR) in cancer. Mechanisms, reversal using 

modulators of MDR and the role of MDR modulators in influencing the pharmacokinetics of anticancer 

drugs, Eur. J. Pharm. Sci., 11, 265–283, 2000. 

111. Juliano, R. L. and Ling, V., A surface glycoprotein modulating drug permeability in Chinese hamster 

ovary cell mutants, Biochim. Acta., 455, 152–162, 1976. 

112. Roe, M. et al., Reversal of P-glycoprotein mediated multidrug resistance by novel anthranilamide 

derivatives, Bioorg. Med. Chem. Lett., 9, 595–600, 1999. 

113. Dantzig, A. H. et al., Reversal of P-glycoprotein-mediated multidrug resistance by a potent cyclopropyldibenzosuberane 

modulator, LY335979, Cancer Res., 56, 4171–4179, 1996. 

114. Wallstab, A. et al., Selective inhibition of MDR1 P-glycoprotein-mediated transport by the acridone 

carboxamide derivative GG918, Br. J. Cancer, 79, 1053–1060, 1999. 

115. Robert, J. and Jarry, C., Multidrug resistance reversal agents, J. Med. Chem., 46, 4805–4817, 

2003. 

116. Thomas, H. and Coley, H. M., Overcoming multidrug resistance in cancer: An update on the clinical 

strategy of inhibiting P-glycoprotein, Cancer Control, 10, 159–165, 2003. 

117. Wandel, C. et al., P-glycoprotein and cytochrome P450 3A inhibition: dissociation of inhibitory 

potencies, Cancer Res., 59, 3944–3948, 1999. 

118. Planting, A. S. T. et al., A phase I and pharmacologic study of the MDR converter GF120918 in 

combination with doxorubicin in patients with advanced solid tumors, Cancer Chemother. Pharmacol., 

55, 91–999, 2005. 

119. Ugazio, E., Cavalli, R., and Gasco, M. R., Incorporation of cyclosporine A in solid lipid nanoparticles 

(SLN), Int. J. Pharm., 241, 341–344, 2002. 

120. Huang, M. et al., Effect of combinations of suboptimal concentrations of P-glycoprotein blockers on 

the proliferation of MDR gene expressing cells, Int. J. Cancer, 65, 389–397, 1996. 

121. Belpomme, D., Gauthier, S., and Pujade-Lauraine, E., Verapamil increases the survival of patients 

with anthracycline-resistant metastatic breast carcinoma, Ann. Oncol., 11, 1471–1476, 2000. 

122. Hyafil, F. et al., In vitro and in vivo reversal of multidrug resistance by GF120918, an acridonecarboxamide 

derivative, Cancer Res., 53, 4595–4602, 1993. 

123. Maliepaard, M. et al., Circumvention of breast cancer resistance protein (BCRP)-mediated resistance 

to camptothecins in vitro using non-substrate drugs or the BCRP inhibitor GF120918, Clin. Cancer 

Res., 7, 935–941, 2001. 

124. den Ouden, D. et al., In vitro effect of GF120918, a novel reversal agent of multidrug resistance, on 

acute leukemia and multiple myeloma cells, Leukemia, 10, 1930–1936, 1996. 

125. Sparreboom, A. et al., Clinical pharmacokinetics of doxorubicin in combination with GF120918, a 

potent inhibitor of MDR1 P-glycoprotein, Anticancer Drugs, 10, 719–728, 1999. 


776 


126. Wong, H. L. et al., Formulation, characterization and in vitro cytotoxicity of solid lipid nanoparticles 

carrying polymer-linked doxorubicin and GG918, a novel P-glycoprotein inhibitor, AAPS J., 6, 

Abstract T3208, 2004. 

127. de Verdiere, A. C. et al., Reversion of multidrug resistance with polyalkylcyanoacrylate nanoparticles: 

Towards a mechanism of action, Br. J. Cancer, 76, 198–205, 1997. 

128. Soma, C. E. et al., Ability of doxorubicin-loaded nanoparticles to overcome multidrug resistance of 

tumor cells after their capture by macrophages, Pharm. Res., 16, 1710–1716, 1999. 

129. Thierry, A. R. et al., Modulation of doxorubicin resistance in multidrug-resistant cells by liposomes, 

FASEB J., 7, 572–579, 1993. 

130. Romsicki, Y. and Sharom, F. J., The membrane lipid environment modulates drug interactions with 

the P-glycoprotein multidrug transporter, Biochemistry, 38, 6887–6896, 1999. 

131. Moore, M. J. and Erlichman, C., Pharmacology of anticancer drugs, In The Basic Science of 

Oncology, Tannock, I. F. and Hill, R. F., Eds., McGraw-Hill, Toronto, pp. 370–391, 1998. 

132. Tabatt, K. et al., Transfection with different colloidal systems: Comparison of solid lipid nanoparticles 

and liposomes, J. Control. Release, 97, 321–332, 2004. 


37 

CONTENTS 

Lipoprotein Nanoparticles 

as Delivery Vehicles for 

Anti-Cancer Agents 

Andras G. Lacko, Maya Nair, and Walter J. McConathy 

37.1 Introduction ....................................................................................................................... 777 

37.2 Lipoproteins as Vehicles for the Targeted Delivery of Anti-Cancer Agents .................. 777 

37.3 LDL as a Delivery Vehicle for Anti-Cancer Agents........................................................ 779 

37.4 HDL as a Delivery Vehicle for Anti-Cancer Agents ....................................................... 779 

37.5 Conclusions ....................................................................................................................... 781 

References.................................................................................................................................... 782 


Nanotechnology and its biomedical applications have exploded in the last several years. Because of 

initiatives in Europe 1 and the United States, 2,3 increased research funding and venture capital 4 have 

been available to facilitate these developments. Nanotechnology, nanoscience, or nanoengineering 

encompass pharmaceutical formulations, devices, chemical processes, or that may be encapsulated 

into a compartment of less than 1,000 nm in diameter. This volume is dedicated to the description of 

a wide variety of drug delivery devices and their application to cancer therapy. 

Although lipoproteins have long been considered as potential vehicles to deliver the “magic 

bullet” for targeted cancer chemotherapy, 5,6 they have yet to be considered as prime candidates for 

clinical applications as recent reviews on targeted drug delivery 7–9 contain only a limited amount of 

discussion on lipoproteins 7 or none at all. 8,9 The purpose of this chapter is to evaluate lipoproteins 

as delivery agents for anti-cancer drugs and to discuss the challenges that have so far prevented 

clinical applications of lipoprotein-based formulations. 

37.2 LIPOPROTEINS AS VEHICLES FOR THE TARGETED DELIVERY 

OF ANTI-CANCER AGENTS 

Plasma lipoproteins are macromolecular complexes composed of specific protein and lipid components. 

The major physiological role of lipoproteins is to transport water-insoluble lipids from 

their point of origin to their respective destinations. Lipoproteins contain an outer shell that is made 

up of phospholipids, apolipoproteins, and unesterified cholesterol (Figure 37.1a) and an interior 

core compartment, accommodating water-insoluble lipids (triacylglycerols and cholesteryl esters). 


777

778 

Phospholipid 

Unesterified cholesterol 

CE CE 

CE TG 

CE 

Apolipoprotein 

Cholesteryl 

Ester 

Triglyceride 

(a) (b) 

Phospholipid 

Unesterified cholesterol 

Apolipoprotein 

Hydrophobic drug 

(paclitaxel) 

FIGURE 37.1 Loading of a hydrophobic drug (paclitaxel) into the core region of a generic lipoprotein 

(Adapted from Wainberg, P. B., Hosp. Pract., 22, 223–227, 1987.) 

This arrangement provides considerable stability to the overall structure of lipoproteins and makes 

them particularly suitable for the transport of hydrophobic drugs (Figure 37.1b). 

Because of the opportunity to convert hydrophilic anti-cancer agents to highly effective hydrophobic 

pro-drugs via chemical modifications, 10–12 there is essentially no structural limit to the type 

and kind of pharmaceutical agents that may be transported by lipoproteins. Plasma lipoproteins are 

spherical particles that range from 5 to over 1,000 nm in diameter (Figure 37.2). Because of their 

superior stability and smaller size, the classes of low density lipoproteins (LDL) and high density 

lipoproteins (HDL) qualify as drug carrying nanoparticles and will be the subject of this review. 

The outstanding features that render lipoproteins suitable carriers of anti-cancer agents may be 

summarized as follows: 

Density (g/ml) 

0.95 

1.006 

1.02 

1.06 

1.10 

1.20 

HDL 

5 

LDL 

IDL 

10 20 40 60 

Diameter (nm) 

FIGURE 37.2 Buoyant density and size distribution of plasma lipoproteins. 



VLDL Chylomicrons 

Chylomicron 

remnants 

80 1,000

Lipoprotein Nanoparticles as Delivery Vehicles for Anti-Cancer Agents 779 

† Lipoproteins represent natural, biologically compatible components providing enhanced 

safety and efficacy. 

† Apolipoproteins provide a signal for the receptor mediated uptake of the whole particle 

by endocytosis 13 or the selective uptake of the drug/core component. 14 

† Lipoproteins nearly perfectly exemplify the “magic bullet” concept 6 via the receptormediated 

uptake of the encapsulated drug. Because of the over expression of lipoprotein 

receptors by malignant cells, 15–17 there is an opportunity to attack malignant cells selectively 

without substantially impacting normal cells. The targeting potential of 

lipoproteins may be further enhanced by attaching ligands to the respective lipid or 

protein components. 

37.3 LDL AS A DELIVERY VEHICLE FOR ANTI-CANCER AGENTS 

The application of LDL as a targeted drug delivery vehicle was initiated by Krieger et al. 18,19 who 

replaced the core of native LDL with cholesteryl linoleate and suggested that hydrophobic 

compounds, including drugs, may be incorporated into LDL for diagnostic or therapeutic purposes 

in a similar manner. 18 Subsequently, Gal et al. 20 proposed an LDL-based drug delivery system for 

treating gynecological neoplasms. To date, numerous studies have been performed with LDL 

particles as targeted drug carriers; nevertheless, the enthusiasm for LDL-based drug delivery 

systems has been modest as indicated by the paucity of approved therapeutic formulations or 

ongoing clinical trials at present. 

The major challenge in the preparation of reconstituted LDL (rLDL) nanoparticles is the 

availability of apolipoprotein B-100 (apoB-100), the major protein component of LDL. 13 In 

addition, the high molecular weight of apoB-100 and its tendency to aggregate upon delipidation 

present additional difficulties in the preparation of rLDL-based drug formulations. Early attempts to 

prepare rLDL/drug complexes were undertaken by Lundberg, 21–23 combining egg yolk phosphatidylcholine 

(EYPC), a hydrophobic drug (cytotoxic mustard carbamate), and apoB-100. 20 The 

resultant nanoparticle had a diameter of about 23 nm and was metabolized by fibroblasts in a 

manner similar to native LDL. 21 A series of studies by Van Berkel et al. developed unilamellar 

liposome-like particles, containing apolipoprotein E that were taken up by tissues via the LDL 

receptor upon injection into rats. 24 Van Berkel et al. also encapsulated a lipophilc derivative of 

daunorubicin into these nanoparticles that were subsequently taken up by B16 tumors 10 via the 

LDL receptor. 25 

Maranhao et al. developed an alternate approach to prepare a drug carrying, protein-free 

microemulsion (LDE) by combining egg yolk phosphatidyl choline, triolein, cholesterol, and 

cholesteryl oleate. 26 The LDE particles apparently associated with apoE upon injection into rats 

and had clearance rates similar to LDL. These LDE particles were subsequently loaded with either 

oleoyl-etoposide 12 or oleoyl-paclitaxel 27 to explore their respective therapeutic potentials. Studies 

with oleoyl-paclitaxel in mice showed a nine-fold increase in LD 50 of oleoyl-paclitaxel encapsulated 

into LDE vs. the free drug. 27 A novel approach for preparing rLDL/drug complexes was 

reported by Owens et al. who utilized peptides representing the receptor binding region of apoB 

instead of full length apo B. 28 This microemulsion resembling LDL was able to support the growth 

of U937 cells similar to native LDL. 29 

37.4 HDL AS A DELIVERY VEHICLE FOR ANTI-CANCER AGENTS 

The utilization of lipoprotein nanoparticles for the delivery of anti-cancer agents is based on the 

hypothesis that rapidly proliferating cells (including cancer cells) have a higher expression of 

lipoprotein receptors 15–17 to meet their increased need for cholesterol (new cell membrane 


780 


synthesis). This view is consistent with the lower plasma cholesterol levels found in cancer 

patients. 30 Clinical studies showed that HDL cholesterol levels were lower in cancer patients 

versus normal controls than their LDL-cholesterol. 31 Similar findings were reported for patients 

with hematologic malignancies, 32–34 small cell lung cancer, 35 and colorectal adenoma. 36 Recent 

findings with the HDL receptors (scavenger receptor B1 [SR-B1], CD36, and LIMPII analogous-1 

[CLA-1]) show that breast cancer cells have substantially higher expression of this HDL receptor 

than the surrounding normal cells. 37 These findings suggest that targeted delivery of anti-cancer 

drugs using an HDL delivery system is likely to be highly effective. 38 

The proliferation of adenocarcinoma 39 and other cancer cells 40,41 were shown to be enhanced 

by HDL or HDL components. In addition, Pussinen et al. have shown that breast cancer cells take 

up cholesterol from HDL via the CLA-1 receptor pathway. 42 These findings show that HDL 

receptors may play a key role in controlling the proliferative capacity of cancer cells, and they 

establish the concept of selective targeting of anti-cancer agents delivered via reconstituted HDL 

(rHDL). These findings are consistent with findings from this laboratory, showing that cancer cells 

produced substantially stronger immunoblots with an SR-BI antibody than normal fibroblasts. 43 It 

has also been shown that the respective uptakes of paclitaxel and cholesteryl esters were highly 

correlated (r 2 Z0.88; pZ0.04) when delivered as core components of rHDL nanoparticles, indicating 

that the paclitaxel was taken up by a receptor-mediated pathway. 43 

The delivery of anti-cancer agents via native HDL is exceedingly challenging because of the 

labor involved in the large scale isolation of human plasma HDL. Concerns about bio-safety would 

likely render the development of the HDL/drug formulations perhaps even more costly than those 

employing LDL. 44 So far, there have been only limited attempts to incorporate drugs into native 

HDL 45 and none with the ultimate purpose of therapeutic applications. Here, rHDL has proven to be 

a much more effective drug delivery vehicle than native HDL. 

The nomenclature currently used in the literature is confusing as to what rHDL actually 

represents. The terms reconstituted and recombinant are both being used to describe artificially 

generated HDL-type nanoparticles although the former designation is becoming more predominant. 

The nomenclature is further confounded by the naming of rHDL nanoparticles based on the 

FIGURE 37.3 Transformation of the discoidal precursor (nascent) HDL to its stable, spherical configuration 

via the accumulation of cholesteryl esters, catalyzed by lecithin: cholesterol acyltransferase (LCAT). (From 

Alexander, E. T., et al., Biochemistry, 44(14), 5409–5419, 2005. With permission.) 



different stages in the maturation of the circulating HDL particle (Figure 37.3). 46,47 HDL metabolism 

commences by the secretion of discoidal, primarily phosphatidyl choline (PC) and 

apolipoprotein A-I (apoA-I) containing particles by the liver 48 and the intestine. These HDL 

precursors are subsequently converted to spherical (mature) HDL 49 by the action of lecithin:cholesterol 

acyltransferase (LCAT) by expanding the particle’s cholesteryl ester core (Figure 37.3). 

Scientific reports and patent disclosures have referred to both discoidal and spherical nanoparticles 

as “rHDL.” 50–54 

For the purpose of formulating drug delivery complexes, spherical nanoparticles have an 

advantage over other configurations as their core region can accommodate the drug to be transported 

(Figure 37.1). Van Berkel et al. prepared spherical neo-high density lipoproteins as drug 

delivery vehicles 54 by incorporating the lipophilic derivative of iododeoxyuridine into neo-HDL, 

resulting in a nanoparticle resembling circulating apoE-free HDL. 11 Earlier work in this laboratory 

yielded stable rHDL/drug nanoparticles composed of Taxol w and the components of circulating 

HDL. 43 The paclitaxel from the rHDL/Taxol w nanoparticles was apparently taken up by cancer 

cells via a receptor mediated mechanism. 43 An enhanced rHDL formulation containing paclitaxel 

has been recently developed with increased drug carrying capacity and cytotoxicity toward several 

cancer cell lines. 55 


Lipoprotein-based nanoparticles have the potential to provide a superior performance compared to 

most alternative formulations for drug delivery agents because of their natural, biodegradable 

ingredients that also help to protect against rapid removal from the circulation by the reticuloendothelial 

system. 44,56 Additional important issues in favor of lipoproteins nanoparticles are the 

biological compatibility of lipoprotein based formulations, 44,56,57 many that have already been 

safely injected into human subjects. 58–62 Lipoprotein nanoparticles have substantially smaller 

molecular diameter than most liposomal drug preparations, a considerable advantage, as even 

smaller liposomes have been considered to have superior pharmacokinetic properties. 63 Another 

important issue favoring lipoprotein nanoparticles over alternate drug delivery systems is their 

targeting potential via receptor-mediated mechanisms that are over expressed in cancer cells vs. 

normal cells. 15,17,37 This type of lipoprotein-based (LDL) drug delivery mechanism has been shown 

to be superior in the delivery of omega-3 fatty acids to cancer cells over serum albumin. 64 

One of the major advantages of lipoprotein nanoparticles in cancer chemotherapy is that their 

targeting potential may be substantially enhanced by attaching ligands to their surface lipid or 

protein components. Targeted drug delivery is of major interest in cancer chemotherapy where 

several monoclonal antibodies have already been employed as targeting agents. 65,66 However, there 

have been only limited attempts to explore this drug delivery model; it has essentially unlimited 

potential in cancer chemotherapy. Earlier attempts to enhance the targeting of the lipoprotein 

nanoparticles was reported by Van Berkel et al. by using lactosylated derivatives of LDL 67,68 and 

neo-HDL. 69 These modified lipoprotein nanoparticles were shown to be preferentially targeted to 

parenchymal liver cells 67,68 that provided a model for converting water soluble drugs to pro-drugs 

for encapsulaton into lipoproteins 69 and a potential drug delivery vehicle for treating hepatitis B. 70 

Additional attempts for enhancing the targeting of lipoprotein nanoparticles involved the attachment 

of monoclonal antibodies (MABs) to LDL to target histocompatibility antigens. 45 Recent 

studies in this laboratory led to the development of an rHDL particle with folate residues attached to 

the apo A-I component. These folate-bearing nanoparticles promoted a substantially enhanced 

uptake of paclitaxel by OVCAR-3 cells from the modified rHDL/paclitaxel complex. 

Despite many potential advantages over other drug delivery models, the application of lipoprotein 

nanoparticles for cancer chemotherapy currently lags behind other drug delivery systems. 

Although lipoprotein nanoparticles are biologically compatible and offer superb targeting potential 


782 

via receptor-mediated mechanisms, the challenge of obtaining the apolipoprotein starting 

material has so far been a major obstacle in the manufacturing of lipoprotein-based 

pharmaceutical formulations. 

In summary, lipoprotein-based nanoparticles have the potential to perform as superior drug 

delivery agents for anti-cancer drugs because of their natural biodegradable components that 

protect against their rapid removal from the circulation and their targeting potential via receptors 

that are over expressed in cancer cells vs. normal cells. 

REFERENCES 


1. European Science Foundation Policy Briefing, ESF scientific forward look on nanomedicine. www. 

esf.org, 2005. 

2. U.S. Department of Health and Human Services, National Institutes of Health, and National Cancer 

Institute, Cancer NANOTECHNOLOGY plan, A strategic initiative to transform clinical oncology 

and basic research through the directed application of nanotechnology, 2004. 

3. NCI Alliance for Nanotechnology in Cancer Awards. http://nano.cancer.gov/alliance_awards/ 

4. Paull, R., Wolfe, J., Hebert, P., and Sinkula, M., Investing in nanotechnology, Nat. Biotechnol., 21, 

1144–1147, 2003. 

5. Counsell, R. E. and Pohland, R. C., Lipoproteins as potential site-specific delivery systems for 

diagnostic and therapeutic agents, J. Med. Chem., 25, 1115–1120, 1982. 

6. Sörgel, F., The return of Ehrlich’s ‘Therapia magna sterilisans’ and other Ehrlich concepts. Series of 

papers honoring Paul Ehrlich on the occasion of his 150th birthday, Chemotherapy, 50, 6–10, 2004. 

7. Dubowchik, G. M. and Walker, M. A., Receptor-mediated and enzyme-dependent targeting of cytotoxic 

anticancer drugs, Pharmacol. Ther., 83, 67–123, 1999. 


1818–1822, 2004. 

9. Guillemard, V. and Saragovi, H. U., Novel approaches for targeted cancer therapy, Curr. Cancer Drug 

Targets, 4, 313–326, 2004. 

10. Versluis, A. J., Rump, E. T., Rensen, P. C., van Berkel, T. J., and Bijsterbosch, M. K., Stable 

incorporation of a lipophilic daunorubicin prodrug into apolipoprotein E-exposing liposomes 

induces uptake of prodrug via low-density lipoprotein receptor in vivo, J. Pharmacol. Exp. Ther., 

289, 1–7, 1999. 

11. de Vrueh, R. L., Rump, E. T., Sliedregt, L. A., Biessen, E. A., van Berkel, T. J., and Bijsterbosch, 

M. K., Synthesis of a lipophilic prodrug of 9-(2-phosphonyl methoxyethyl)adenine (PMEA) and its 

incorporation into a hepatocyte-specific lipidic carrier, Pharm. Res., 16, 1179–1185, 1999. 

12. Azevedo, C. H., Carvalho, J. P., Valduga, C. J., and Maranhao, R. C., Plasma kinetics and uptake by 

the tumor of a cholesterol-rich microemulsion (LDE) associated to etoposide oleate in patients with 

ovarian carcinoma, Gynecol. Oncol., 97, 178–182, 2005. 

13. Segrest, J. P., Jones, M. K., De Loof, H., and Dashti, N., Structure of apolipoprotein B-100 in low 

density lipoproteins, J. Lipid Res., 42, 1346–1367, 2001. 

14. Williams, D. L., Temel, R. E., and Connelly, M. A., Roles of scavenger receptor BI and APO A-I in 

selective uptake of HDL cholesterol by adrenal cells, Endocr. Res., 26, 639–651, 2000. 

15. Tosi, M. R. and Tugnoli, V., Cholesteryl esters in malignancy, Clin. Chim. Acta, 359, 27–45, 2005. 

16. Ho, Y. K., Smith, R. G., Brown, M. S., and Goldstein, J. L., Low-density lipoprotein (LDL) receptor 

activity in human acute myelogenous leukemia cells, Blood, 52, 1099–1114, 1978. 

17. Imachi, H., Murao, K., Sayo, Y., Hosokawa, H., Sato, M., Niimi, M., Kobayashi, S., Miyauchi, A., 

Ishida, T., and Takahara, J., Evidence for a potential role for HDL as an important source of cholesterol 

in human adrenocortical tumors via the CLA-1 pathway, Endocr. J., 46, 27–34, 1999. 

18. Krieger, M., Brown, M. S., Faust, J. R., and Goldstein, J. L., Replacement of endogenous cholesteryl 

esters of low density lipoprotein with exogenous cholesteryl linoleate. Reconstitution of a biologically 

active lipoprotein particle, J. Biol. Chem., 253, 4093–4101, 1978. 

19. Krieger, M., Goldstein, J. L., and Brown, M. S., Receptor-mediated uptake of low density lipoprotein 

reconstituted with 25-hydroxycholesteryl oleate suppresses 3-hydroxy-3-methylglutaryl-coenzyme. A 

reductase and inhibits growth of human fibroblasts, Proc. Natl Acad Sci. U.S.A. 75, 5052–5056, 1978. 



20. Gal, D., Ohashi, M., MacDonald, P. C., Buchsbaum, H. J., and Simpson, E. R., Low-density lipoprotein 

as a potential vehicle for chemotherapeutic agents and radionucleotides in the management of 

gynecologic neoplasms, Am. J. Obstet. Gynecol., 139, 877–885, 1981. 

21. Lundberg, B., Preparation of drug-low density lipoprotein complexes for delivery of antitumoral 

drugs via the low density lipoprotein pathway, Cancer Res., 47, 4105–4108, 1987. 

22. Lundberg, B. and Suominen, L., Preparation of biologically active analogs of serum low density 

lipoprotein, J. Lipid Res., 25, 550–558, 1984. 

23. Lundberg, B., Cytotoxic activity of two new lipophilic steroid nitrogen carbamates incorporated into 

low-density lipoprotein, Anticancer Drug Des., 9, 471–476, 1994. 

24. Rensen, P. C., Schiffelers, R. M., Versluis, A. J., Bijsterbosch, M. K., Van Kuijk-Meuwissen, M. E., 

and Van Berkel, T. J., Human recombinant apolipoprotein E-enriched liposomes can mimic lowdensity 

lipoproteins as carriers for the site-specific delivery of antitumor agents, Mol. Pharmacol., 52, 

445–455, 1997. 

25. Versluis, A. J., Rensen, P. C., Rump, E. T., Van Berkel, T. J., and Bijsterbosch, M. K., Low-density 

lipoprotein receptor-mediated delivery of a lipophilic daunorubicin derivative to B16 tumours in mice 

using apolipoprotein E-enriched liposomes, Br. J. Cancer, 78, 1607–1614, 1998. 

26. Maranhão, R. C., Cesar, T. B., Pedroso-Mariani, S. R., Hirata, M. H., and Mesquita, C. H., Metabolic 

behavior in rats of a nonprotein microemulsion resembling low-density lipoprotein, Lipids, 28, 

691–696, 1993. 

27. Rodrigues, D. G., Maria, D. A., Fernandes, D. C., Valduga, C. J., Couto, R. D., Ibañez, O. C., and 

Maranhão, R. C., Improvement of paclitaxel therapeutic index by derivatization and association to a 

cholesterol-rich microemulsion: In vitro and in vivo studies, Cancer Chemother. Pharmacol., 55, 

565–576, 2005. 

28. Owens, M. D., Baillie, G., and Halbert, G. W., Physicochemical properties of microemulsion analogues 

of low density lipoprotein containing amphiphatic apoprotein B receptor sequences, Int. 

J. Pharm., 228, 109–117, 2001. 

29. Baillie, G., Owens, M. D., and Halbert, G. W., A synthetic low density lipoprotein particle capable of 

supporting U937 proliferation in vitro, J. Lipid Res., 43, 69–73, 2002. 

30. Markel, A. and Brook, G. J., Cancer and hypocholesterolemia, Isr. J. Med. Sci., 30, 787–793, 1994. 

31. Fiorenza, A. M., Branchi, A., and Sommariva, D., Serum lipoprotein profile in patients with cancer. 

A comparison with non-cancer subjects, Int. J. Clin. Lab Res., 30, 141–145, 2000. 

32. Dessì, S., Batetta, B., Pulisci, D., Accogli, P., Pani, P., and Broccia, G., Total and HDL cholesterol in 

human hematologic neoplasms, Int. J. Hematol., 54, 483–486, 1991. 

33. Moschovi, M., Trimis, G., Apostolakou, F., Papassotiriou, I., and Tzortzatou-Stathopoulou, F. J., 

Serum lipid alterations in acute lymphoblastic leukemia of childhood, Pediatr. Hematol. Oncol., 

26, 289–293, 2004. 

34. Scribano, D., Baroni, S., Pagano, L., Zuppi, C., Leone, G., and Giardina, B., Prognostic relevance of 

lipoprotein cholesterol levels in acute lymphocytic and nonlymphocytic leukemia, Haematologica, 

81, 343–345, 1996. 

35. Siemianowicz, K., Gminski, J., Stajszczyk, M., Wojakowski, W., Goss, M., Machalski, M., Telega, 

A., Brulinski, K., and Magiera-Molendowska, H., Serum HDL cholesterol concentration in patients 

with squamous cell and small cell lung cancer, Int. J. Mol. Med., 6, 307–311, 2000. 

36. Bayerdörffer, E., Mannes, G. A., Richter, W. O., Ochsenkühn, T., Seeholzer, G., Köpcke, W., 

Wiebecke, B., and Paumgartner, G., Decreased high-density lipoprotein cholesterol and increased 

low-density cholesterol levels in patients with colorectal adenomas, Ann. Int. Med., 118, 481–487, 

1993. 

37. Cao, W. M., Murao, K., Imachi, H., Yu, H., Abe, H., Yamauchi, A., Niimi, M., Miyauchi, A., Wong, 

N. C., and Ishida, T., A mutant high-density lipoprotein receptor inhibits proliferation of human breast 

cancer cells, Cancer Res., 64, 1515–1521, 2004. 

38. Lacko, A. G., Nair, M. P., Paranjape, S., Johnson, S., and McConathy, W. J., A novel delivery system 

for targeted cancer therapy, In Stem Cell and Targeted Therapy, Dickey, K. A. and Keating, A., Eds., 

Garden Jennings, Charlottesville, VA, pp. 179–183, 2003. 

39. Favre, G., Tazi, K. A., Le Gaillard, F., Bennis, F., Hachem, H., and Soula, G., High density lipoprotein3 

binding sites are related to DNA biosynthesis in the adenocarcinoma cell line A549, J. Lipid 

Res., 34, 1093–1106, 1993. 


784 


40. Gospodarowicz, D., Lui, G. M., and Gonzalez, R., High-density lipoproteins and the proliferation of 

human tumor cells maintained on extracellular matrix-coated dishes and exposed to defined medium, 

Cancer Res., 42, 3704–3713, 1982. 

41. Jozan, S., Faye, J. C., Tournier, J. F., Tauber, J. P., David, J. F., and Bayard, F., Interaction of estradiol 

and high density lipoproteins on proliferation of the human breast cancer cell line MCF-7 adapted to 

grow in serum free conditions, Biochem. Biophys. Res. Commun., 133, 105–112, 1985. 

42. Pussinen, P. J., Karten, B., Wintersperger, A., Reicher, H., McLean, M., Malle, E., and Sattler, W., 

The human breast carcinoma cell line HBL-100 acquires exogenous cholesterol from high-density 

lipoprotein via CLA-1 (CD-36 and LIMPII analogous 1)-mediated selective cholesteryl ester uptake, 

Biochem. J., 349, 559–566, 2000. 

43. Lacko, A. G., Nair, M., Paranjape, S., Johnson, S., and McConathy, W. J., High density lipoprotein 

complexes as delivery vehicles for anticancer drugs, Anticancer Res., 22, 2045–2049, 2002. 

44. De Smidt, P. C. and van Berkel, T. J., LDL-mediated drug targeting, Crit. Rev. Ther. Drug Carrier. 

Syst., 7, 99–120, 1990. 

45. Shaw, J. M. and Shaw, K. V., Key issues in the delivery of pharmacological agents using lipoproteins: 

Design of a synthetic apoprotein-lipid carrier, Targeted Diagn. Ther., 5, 351–383, 1991. 

46. Barter, P. J., Hugh sinclair lecture: The regulation and remodelling of HDL by plasma factors, 

Atheroscler. Suppl., 3, 39–47, 2002. 

47. Linsel-Nitschke, P. and Tall, A. R., HDL as a target in the treatment of atherosclerotic cardiovascular 

disease, Nat. Rev. Drug Discov., 4, 193–205, 2005. 

48. Hamilton, R. L., Williams, M. C., Fielding, C. J., and Havel, R. J., Discoidal bilayer structure of 

nascent high density lipoproteins from perfused rat liver, J. Clin. Invest., 58, 667–680, 1976. 

49. Forte, T., Norum, K. R., Glomset, J. A., and Nichols, A. V., Plasma lipoproteins in familial lecithin: 

Cholesterol acyltransferase deficiency: Structure of low and high density lipoproteins as revealed by 

electron microscopy, J. Clin. Invest., 50, 1141–1148, 1971. 

50. Lerch, P. G., Fortsch, V., Hodler, G., and Bolli, R., Production and characterization of a reconstituted 

high density lipoprotein for therapeutic applications, Vox Sang., 71, 155–164, 1996. 

51. Levine, D. M., Simon, S. R., Gordon, B. R., Parker, T. S., Saal, S. D., and Rubin, A. L., United States 

Patent no. 5,128,318. 

52. Sparks, D. L., Phillips, M. C., and Lund-Katz, S., The conformation of apolipoprotein A-I in discoidal 

and spherical recombinant high density lipoprotein particles 13C. NMR studies of lysine ionization 

behavior, J. Biol. Chem., 267, 25830–25838, 1992. 

53. Jonas, A., Wald, J. H., Toohill, K. L., Krul, E. S., and Kézdy, K. E., The conformation of apolipoprotein 

A-I in discoidal and spherical recombinant high density lipoprotein particles. 13C NMR 

studies of lysine ionization behavior, J. Biol. Chem., 265, 22123–22129, 1990. 

54. Schouten, D., van der Kooij, M., Muller, J., Pieters, M. N., Bijsterbosch, M. K., and van Berkel, T. J., 

Development of lipoprotein-like lipid particles for drug targeting: Neo-high density lipoproteins, Mol. 

Pharmacol., 44, 486–492, 1993. 

55. Lacko, A. G., Nair, M., Paranjape, S., Mooberry, L., and McConathy, W. J., Advanced Drug Delivery 

System for Breast Cancer Chemotherapy Department of Defense, Congressionally Directed Medical 

Research Program “ERA of Hope” Breast Cancer Research Review, Philadelphia, Pennsylvania, 

2005. http://mrmcweb4.detrick.army.mil/bcrp/era/abstracts2005/0110582_abs.pdf 

56. Shaw, J. M., Shaw, K. V., Yanovich, S., Iwanik, M., Futch, W. S., Rosowsky, A., and Schook, L. B., 

Delivery of lipophilic drugs using lipoproteins, Ann. NY Acad. Sci., 507, 252–271, 1987. 

57. Adams, T., Alanazi, F., and Lu, D. R., Safety and utilization of blood components as therapeutic 

delivery systems, Curr. Pharm. Biotechnol., 4, 275–282, 2003. 

58. Nanjee, M. N., Crouse, J. R., King, J. M., Hovorka, R., Rees, S. E., Carson, E. R., Morgenthaler, J. J., 

Lerch, P., and Miller, N. E., Effects of intravenous infusion of lipid-free apo A-I in humans, Arterioscler. 

Thromb. Vasc. Biol., 16, 1203–1214, 1996. 

59. Nanjee, M. N., Doran, J. E., Lerch, P. G., and Miller, N. E., Acute effects of intravenous infusion of 

ApoA1/phosphatidylcholine discs on plasma lipoproteins in humans, Arterioscler. Thromb. Vasc. 

Biol., 19, 979–989, 1999. 

60. Bisoendial, R. J., Hovingh, G. K., Levels, J. H., Lerch, P. G., Andresen, I., Hayden, M. R., Kastelein, 

J. J., and Stroes, E. S., Restoration of endothelial function by increasing high-density lipoprotein in 

subjects with isolated low high-density lipoprotein, Circulation, 107, 2944–2948, 2003. 



61. Pajkrt, D., Doran, J. E., Koster, F., Lerch, P. G., Arnet, B., van der Poll, T., ten Cate, J. W., and van 

Deventer, S. J., Antiinflammatory effects of reconstituted high-density lipoprotein during human 

endotoxemia, J. Exp. Med., 184, 1601–1608, 1996. 

62. Nanjee, M. N., Cooke, C. J., Garvin, R., Semeria, F., Lewis, G., Olszewski, W. L., and Miller, N. E., 

J. Lipid Res., 42, 1586–1593, 2001. 

63. Palatini, P., Disposition kinetics of phospholipid liposomes, Adv. Exp. Med. Biol., 318, 375–391, 1992. 

64. Edwards, I. J., Berquin, I. M., Sun, H., O’Flaherty, J. T., Daniel, L. W., Thomas, M. J., Rudel, L. L., 

Wykle, R. L., and Chen, Y. Q., Differential effects of delivery of omega-3 fatty acids to human cancer 

cells by low-density lipoproteins versus albumin, Clin. Cancer Res., 10, 8275–8283, 2004. 

65. Toi, M., Takada, M., Bando, H., Toyama, K., Yamashiro, H., Horiguchi, S., and Saji, S., Current 

status of antibody therapy for breast cancer, Breast Cancer, 11, 10–14, 2004. 

66. Boskovitz, A., Wikstrand, C. J., Kuan, C. T., Zalutsky, M. R., Reardon, D. A., and Bigner, D. D., 

Monoclonal antibodies for brain tumour treatment, Expert. Opin. Biol. Ther., 4, 1453–1471, 2004. 

67. Bijsterbosch, M. K. and Van Berkel, T. J., Uptake of lactosylated low-density lipoprotein by galactose-specific 

receptors in rat liver, Biochem. J., 270, 233–239, 1990. 

68. Bijsterbosch, M. K. and Van Berkel, T. J., Lactosylated high density lipoprotein: A potential carrier 

for the site-specific delivery of drugs to parenchymal liver cells, Mol. Pharmacol., 41, 404–411, 1992. 

69. de Vrueh, R. L., Rump, E. T., Sliedregt, L. A., Biessen, E. A., van Berkel, T. T., and Bijsterbosch, 

M. K., Synthesis of a lipophilic prodrug of 9-(2-phosphonyl methoxyethyl)adenine (PMEA) and its 

incorporation into a hepatocyte-specific lipidic carrier, Pharm. Res., 16, 1179–1185, 1999. 

70. de Vrueh, R. L., Rump, E. T., van De Bilt, E., van Veghel, R., Balzarini, J., Biessen, E. A., van Berkel, 

T. J., and Bijsterbosch, M. K., Carrier-mediated delivery of 9-(2-phosphonylmethoxyethyl)adenine to 

parenchymal liver cells: A novel therapeutic approach for hepatitis B, Antimicrob. Agents 

Chemother., 44, 477–483, 2000. 


38 

CONTENTS 

DQAsomes as Mitochondria- 

Targeted Nanocarriers for 

Anti-Cancer Drugs 

Shing-Ming Cheng, Sarathi V. Boddapati, Gerard 

G. M. D’Souza, and Volkmar Weissig 

38.1 Introduction ....................................................................................................................... 787 

38.2 Apoptosis and Mitochondria ............................................................................................. 788 

38.3 Proapoptotic Drugs Acting on Mitochondria ................................................................... 788 

38.4 Mitochondriotropic Vesicles (DQAsomes)....................................................................... 790 

38.5 DQAsome-Mediated Delivery of pDNA to Mitochondria in Living 

Mammalian Cells .............................................................................................................. 791 

38.6 Encapsulation of Paclitaxel into DQAsomes.................................................................... 793 

38.7 DQAsomal-Encapsulated Paclitaxel Triggers Apoptosis In Vitro ................................... 794 

38.8 Tumor Growth Inhibition Study with Paclitaxel-Loaded DQAsomes In Vivo ............... 796 


References..................................................................................................................................... 797 


A major challenge in treating cancer is the lack of selectivity of the currently available cytotoxic 

agents. Many such agents fail to significantly distinguish between cancer cells and healthy cells. 

Consequently, systemic application of these drugs is prone to cause severe side effects in other 

tissues, greatly limiting the maximal allowable dose of the drug. Another major hurdle to successfully 

treating cancer is multi-drug resistance (MDR). There are multiple putative origins of MDR, 

but one of the most studied today is the efflux pump that serves to remove drug molecules from the 

cell before they can act at their particular subcellular target inside the cell. Both problems, i.e., 

non-specificity and MDR, could potentially be solved by the development of highly selective 

tumor-specific drug delivery systems. Progress in this direction has already been made with the 

development and subsequent Food and Drug Administration (FDA) approval of Doxil w , the first 

liposomal anti-cancer drug, and similar systems are under development. However, the DQAsomebased 

approach described in this chapter takes drug delivery one significant step farther. To 

efficiently and selectively eradicate carcinoma cells, the cytotoxic drug not only needs to be 

delivered to the tumor cell, but it must also be delivered to the particular target inside the cell. 

Tumor-specific subcellular drug delivery constitutes a new approach for the chemotherapy 

of cancer. This approach becomes increasingly feasible as the particular mechanism of action, 


787

788 

i.e., the molecular targets of anti-cancer drugs, is being recognized. Transporting the cytotoxic drug 

to its intracellular target could potentially overcome MDR by bypassing the p-glycoprotein, i.e., the 

drug would literally be hidden from the p-glycoprotein inside the delivery system until it becomes 

selectively released at the particular intracellular site of action. At the same time, a sub-cellular 

delivery system would significantly increase the subcellular bioavailability of any drug acting 

inside a cell. 

38.2 APOPTOSIS AND MITOCHONDRIA 

Apoptosis (programmed cell death) plays a central role in tissue homeostasis, and it is generally 

recognized that inhibition of apoptosis may contribute to cell transformation. 1 Significant knowledge 

has been accumulated during the past several years about how the apoptotic machinery is 

controlled. 2–20 As each new regulatory mechanism had been identified, dysfunction of that 

mechanism has been linked to one or another type of cancer. 17 Dysregulation of the apoptotic 

machinery is now generally accepted as an almost universal component of the transformation 

process of normal cells into cancer cells. 

A large body of experimental data demonstrates that mitochondria play a key role in the 

complex apoptotic mechanism. 18–49 Mitochondria have been shown to trigger cell death via 

several mechanisms: by disrupting electron transport and energy metabolism, by releasing or 

activating proteins that mediate apoptosis, and by altering cellular redox potential. A critical 

event leading to programmed cell death is the mitochondrial membrane permeabilization that is 

under the control of the permeability transition pore complex (mPTPC), a multiprotein complex 

formed at the contact site between the mitochondrial inner and outer membranes. The mPTPC is 

widely accepted as being central to the process of cell death and has accordingly been recommended 

as a privileged pharmacological target for cytoprotective and for cytotoxic therapies 

in general. 50 In particular, it has been suggested that targeting specific mPTPC components may 

overcome bcl-2 mediated apoptosis inhibition in cancer cells. 1 Several studies have already demonstrated 

the feasibility of eliminating neoplastic cells by selectively inducing apoptosis (reviewed in 

reference 17). The design of mitochondria-targeted cytotoxic drugs has been formulated as a novel 

strategy for overcoming apoptosis resistance in tumor cells, 1 which, intriguingly, opens up a whole 

new avenue for the therapy of cancer by “tricking cancer cells into committing suicide.” 17 

38.3 PROAPOPTOTIC DRUGS ACTING ON MITOCHONDRIA 


Several conventional anti-cancer drugs, such as doxorubicin, and cisplatin, have no direct effect on 

mitochondria. 51 These conventional chemotherapeutic agents elicit mitochondrial permeabilization 

in an indirect fashion by induction of endogenous effectors that are involved in the physiologic 

control of apoptosis. 1 However, a variety of clinically approved drugs such as paclitaxel, 52–60 

VP-16 (etoposide) 61–64 and vinorelbine 58 as well as an increasing number of experimental anticancer 

drugs such as betulinic acid, lonidamine, CD-437 (a synthetic retinoids) and ceramide 

(reviewed in 1 ) have been found to act directly on mitochondria resulting in triggering apoptosis. 

These agents may induce apoptosis in circumstances in which conventional drugs fail to act 

because endogenous apoptosis inducing pathways, e.g., such as those involving p53, death 

receptors or apical caspase activation, are disrupted, leading to the apoptosis-resistance of tumor 

cells. For example, several in vitro and in vivo studies have shown that the synthetic retinoid CD437 

is able to induce apoptosis in human lung, breast, cervical and ovarian carcinoma cells (reviewed in 

Kaufmann and Gores 2000). It could be demonstrated that in intact cells, CD437-dependent caspase 

activation is preceded by the release of cytochrome C from mitochondria. 65 Moreover, it was 

shown that when added to isolated mitochondria, CD437 causes membrane permeabilization and 

that this effect is prevented by inhibitors of the mPTPC such as cyclosporine A. CD437 constitutes 


DQAsomes as Mitochondria-Targeted Nanocarriers for Anti-Cancer Drugs 789 

an experimental drug that exerts its cytotoxic effect via the mPTPC, i.e., by acting directly at the 

surface or inside of mitochondria. 

The development of anti-cancer drugs whose cytotoxic effects depend on their direct interaction 

with mitochondria inside of living cells raises the issue of the intracellular distribution of these 

drugs after having been taken up by the cell, i.e., the question of their intracellular bioavailability. 

Independent of their mode of cell entry that should mostly take place via passive diffusion through 

the cell membrane, drug molecules become randomly distributed among all cell organelles and, 

depending on the chemical nature of the drug, eventually metabolized. Therefore, anti-cancer drugs 

that exert their cytotoxic activity by directly acting at or inside of mitochondria in living cells would 

dramatically benefit from a delivery system that can selectively transport these drugs to and 

into mitochondria. 

One of the major roles of mitochondria in the metabolism of eukaryotic cells is the synthesis of 

Adenosine triphosphate (ATP) by oxidative phosphorylation via the respiratory chain. According to 

Mitchell’s chemiosmotic hypothesis, electrons from the hydrogens on Nicotine amide adenine 

dinucleotide (NADH) and Flavin adenine dinucleotide (FADH 2) are carried along the respiratory 

chain at the mitochondrial inner membrane, thereby releasing energy that is used to pump protons 

across the inner membrane from the mitochondrial matrix into the intermembrane space. This 

process creates a transmembrane electrochemical gradient that includes contributions from both 

a membrane potential (negative inside) and a pH difference (acidic outside). The membrane 

potential of mitochondria in vitro is between 180 and 200 mV, the maximum a lipid bilayer can 

sustain while maintaining its integrity. 66 Although this potential is reduced in living cells and 

organism to about 130–150 mV as a result of metabolic processes such as ATP synthesis and 

ion transport, 67 it is by far the largest within cells. 

Most interestingly, carcinoma cells posses a different mitochondrial membrane potential 

relative to normal cells. It has been found that in many carcinoma cell lines, the mitochondrial 

membrane potential is higher than in normal epithelial cells (reviewed in reference 68). For 

example, the difference of the mitochondrial membrane potential between the colon carcinoma 

cell line CX-1 and the control green monkey kidney epithelial cell line CV-1 has been reported to 

be approximately 60 mV. 68 Moreover, some carcinoma cells, in particular, human breast adenocarcinoma-derived 

cells, have in addition to the higher mitochondrial membrane potential also an 

elevated plasma membrane potential relative to their normal parent cell lines. 68–75 

The striking difference between normal cells and human adenocarcinoma cells regarding the 

electrical charge of both plasma and mitochondrial membranes has lead numerous investigators 

during the 1990s to explore fundamentally new strategies for the selective targeting of cancer cells. 

Their attempts have been based on compounds with a delocalized charge that have long been 

known to accumulate in mitochondria of living cells in response to the mitochondrial membrane 

potential. Many of these DLCs are toxic to mitochondria at high concentration. For example, 

the rhodacyanine official name of this compound, unknown what the letters stand for 

(MKT-077 71,76–80 ) was the first DLC to be approved by the FDA for clinical trials for the treatment 

of carcinoma. 81 The trials were discontinued, however, because efficacy in tumor cell killing was 

not demonstrated at the particular approved dosage and drug regimen. 68 From the phase I clinical 

trial, it was concluded that it is feasible to target carcinoma cell mitochondria with rhodacyanine 

analogues if drugs with higher therapeutic indices could be developed. 81 

The DQAsomal-based strategy involving the use of dequalinium chloride, a typical representative 

of DLCs, for tumor and mitochondria-specific targeting factually starts where these failed 

attempts of the 1990s have left off. This approach combines the well-proven ability of DLCs to 

specifically target carcinoma cell mitochondria with the selective delivery of apoptotically active 

compounds known to trigger programmed cell death via directly acting at the mitochondrial 

surface. The DQAsomal approach utilizes the mitochondria-specific affinity of DLCs for the 

delivery of pro-apoptotic drugs to mitochondria. Therefore, the actual cytotoxic effect is caused 


790 

by apoptosis-triggering drugs. Any inherent cytotoxicity of the carrier system, i.e., the DLCs used, 

might only add to the overall efficiency of killing carcinoma cells. 

38.4 MITOCHONDRIOTROPIC VESICLES (DQASOMES) 

Dequalinium chloride (DQA, Figure 38.1a) represents a typical mitochondriotropic delocalized 

cation that already almost 20 years ago was shown to selectively accumulate in carcinoma cell 

mitochondria. 73 Most interestingly in the context of this chapter, it was demonstrated recently that 

dequalinium B, a new boron carrier for neutron capture therapy, also accumulates preferentially in 

carcinoma cells over non-transformed cells. 82 

Dequalinium is a dicationic compound resembling “bola”-form electrolytes, i.e., it is a symmetrical 

molecule with two charge centers separated at a relatively large distance. Such symmetric 

bola-like structures are well known from archaeal lipids that usually consist of two glycerol 

backbones connected by two hydrophobic chains. 85,86 The self-assembly behavior of bipolar 

lipid from Archaea has been extensively studied (reviewed in Gambacorta, Gliozi, and De Rosa 

1995). Generally, it has been shown that these symmetric bipolar archaeal lipids can self-associate 

into mechanically, very stable monolayer membranes. The most striking structural difference 

between dequalinium and archaeal lipids lies in the number of bridging hydrophobic chains 

between the polar head groups. Contrary to the common arachaeal lipids, in dequalinium there 

is only one carbohydrate chain that connects the two cationic hydrophilic head groups. Therefore, 

this type of bola lipids has been named single-chain bola-amphiphile. 87,88 The self-association 

behavior of this single-chain cationic bola amphiphile was investigated using Monte Carlo 

computer simulations (Figure 38.1c, left panel), 83 several electron microscopic (EM) techniques 

(Figure 38.1c, right panel) as well as dynamic laser light scattering. 84 

It was found that, upon sonication, dequalinium forms spherical aggregates with diameters 

between about 70 and 700 nm that were termed DQAsomes. 84 Freeze fracture images 

(Figure 38.1) show both convex and concave fracture faces. These images strongly indicate the 

(a) 

(b) 

(c) 

NH 2 

CH 3 

0.2μm 

OR 

CH 3 

N CH CH CH CH CH CH CH CH CH 2 2 2 2 2 2 2 2 CH N 

2 


NH 2 

A B C 

FIGURE 38.1 (a) Chemical structure of dequalinium chloride with overlaid colors indicating, in blue, the 

hydrophilic part and in yellow the hydrophobic part of the molecule. (b) Theoretical possible conformations of 

dequalinium chloride, i.e., stretched versus horseshoe conformation, leading to either a monolayer or a bilayer 

membranous structure following the process of self-assembly. (c) Left panel: Monte Carlo Computer 

Simulations demonstrate the possible self-assembly of dequalinium chloride into vesicles. (From Weissig, 

V., Mogel, H. J., Wahab, M., and Lasch, J., Proceed. Intl. Symp. Control. Rel. Bioact. Mater., 25, 312, 1998.) 

Right panel: Electron microscopic images of vesicles (DQAsomes) made from dequalinium chloride, from left 

to right: Negatively stained, rotary shadowed, freeze fractured. (From Weissig, V., Lasch, L., Erdos, G., 

Meyer, H. W., Rowe, T. C., and Hughes, J., Pharm. Res., 15(2), 334–337, 1998. With permission.) 



FIGURE 38.2 Top panel: Structure of the cyclohexyl derivative of dequalinium. Bottom panel: Schematic 

illustration of the stabilizing effect of the cyclohexyl ring system (black circles). 

liposome-like aggregation of dequalinium. Negatively stained samples (Figure 38.1) demonstrate 

that the vesicle is impervious to the stain and appears as a clear area surrounded by stain with no 

substructure visible. Particle size measurements of DQAsomes stored at room temperature for 24 and 

96 h (not shown) do not show any significant changes in their size distribution in comparison to 

freshly made vesicles measured after one hour. 84 This indicates that DQAsomes do not seem to 

precipitate, to fuse with each other, or to aggregate in solution over a period of several days. The use 

of dequalinium derivatives for the preparation of DQAsome-like vesicles leads to vesicles with 

different size distributions. 89 For example, substituting the methyl group by an aliphatic ring 

system (Figure 38.2) confers superior vesicle forming properties to this bolaamphiphile. These 

DQAsome-like vesicles, i.e., vesicles prepared from the cyclohexyl-derivative of DQA shown in 

Figure 38.2, have a very narrow size distribution of 169G50 nm and can be stored at room temperature 

for at least five months. 

In contrast to vesicles made from dequalinium, bolasomes made from the cyclohexyl derivative 

are also stable upon dilution of the original vesicle preparation. Whereas dequalinium-based bolasomes 

upon dilution slowly disintegrate over a period of several hours, bolasomes made from the 

cyclohexyl compound do not show any change in size distribution following dilution. It appears 

that bulky aliphatic residues attached to the quinolinium heterocycle favor self-association of the 

planar ring system. It has, therefore, been speculated that the bulky group sterically prevents the 

free rotation of the hydrophilic head of the amphiphile around the CH2-axis (Figure 38.2, bottom 

panel), contributing to improved intermolecular interactions between the amphiphilic monomers. 

38.5 DQASOME-MEDIATED DELIVERY OF pDNA TO MITOCHONDRIA 

IN LIVING MAMMALIAN CELLS 

During efforts in developing a mitochondria-specific DNA delivery vector, it could be demonstrated 

that DQAsomes are able to selectively deliver plasmid DNA (pDNA) to mitochondria 

within living mammalian cells. It has been shown, in particular, that DQAsomes stably incorporate 

pDNA, 84 protect the pDNA from nuclease digestion, and mediate its cellular uptake most likely via 

non-specific endocytosis. 90 Using membrane-mimicking liposomal membranes and isolated rat 


792 


liver mitochondria, it was shown that DQAsome/pDNA complexes become destabilized upon 

contact with mitochondrial membranes, but not at cell plasma membranes. 91,92 

The selective destabilization of DQAsomes at mitochondrial membranes may either be caused 

by the difference in the lipid composition between cytoplasmic and mitochondrial membranes 91 or 

by the membrane potential-driven diffusion of individual dequalinium molecules from the 

DQAsome/pDNA complex into the mitochondrial matrix leading to the disintegration of the 

complex. However, the fact that DQAsomes do not lose their cargo (pDNA) during their contact 

with the plasma cell membrane (i.e., during their cellular uptake) but do release the entrapped 

pDNA upon contact with mitochondria appears as a very attractive feature of DQAsomes as a 

mitochondria-specific drug delivery system. Any encapsulated drug would potentially be released 

upon contact with mitochondrial membranes leading to the desired high local drug concentration in 

the immediate proximity to the mPTPC that is recognized as a major target for apoptotically active 

experimental drugs. 1,93,94 

Before translocating to mitochondria in response to the mitochondrial membrane potential, 

however, endocytosed DQAsomes have to be released from endosomes into the cytosol. From 

studies about the intracellular fate of cationic liposome/DNA complexes (lipoplexes), it is known 

that cationic lipids exert a destabilizing effect on endosomal membranes, leading to the release of at 

least a fraction of the lipoplex from early endosomes. 95–98 In agreement with these data, it was 

found that dequalinium-based DQAsomes also display endosomolytic activity. 99 It was shown that 

adding DQAsomes to liposomes mimicking the lipid composition of endosomal membranes reproducibly 

lead to the release of liposomal encapsulated fluorescence marker. 99 

Direct evidence for the ability of DQAsomes to transport pDNA selectively to the site of 

mitochondria was provided by studying the intracellular distribution of mitochondrial leader 

sequence peptide -pDNA conjugates in cultured BT20 cells using confocal fluorescence 

microscopy. 100 Figure 38.3 shows images representative of the confocal fluorescence micrographs 

obtained in this study. The green and red channels used to generate the overlaid images in the far 

right column are shown separately in the preceding columns to facilitate a careful comparison of the 

observed staining patterns. 

The characteristic mitochondrial staining pattern seen with the red channel is a strong indicator 

of mitochondrial viability in the imaged cells. From the composite image obtained by overlaying 

the green and red channels, it can be seen that a sizeable fraction of the intracellular green 

fluorescence co-localized with the red mitochondrial fluorescence (depicted as white areas in 

Figure 38.3c). These observations indicate that, in addition to mediating the cellular uptake of 

the pDNA conjugate, the use of DQAsomes resulted in a definite association of an appreciable 

fraction of the internalized conjugate with mitochondria. 

FIGURE 38.3 (See color insert following page 522.) Representative confocal fluorescence micrographs of 

BT20 cells stained with Mitotrackerw Red CMXRos (red) after exposure to fluorescein labeled linearized 

MLS-pDNA conjugate (green) complexed with DQAsomes; (a) red channel, (b) green channel, (c) overlay of 

red and green channels with white, indicating co-localization of red and green fluorescence.(From D’Souza, G. 

G., Boddapati, S. V., and Weissig, V., Mitochondrion, 5(5), 352–358, 2005. With permission.) 



38.6 ENCAPSULATION OF PACLITAXEL INTO DQASOMES 

Paclitaxel, generally known as an anti-microtubule agent, has recently been demonstrated to trigger 

apoptosis by directly acting on mitochondria. 54,55,58 It has been shown that clinically relevant 

concentrations of paclitaxel directly target mitochondria and trigger apoptosis by inducing 

cytochrome c (cyt c) release in a permeability transition pore (PTP)-dependent manner. 54 This 

mechanism of action is known from other pro-apoptotic, directly on mitochondria acting agents. 51 

A 24-hour delay between the treatment with paclitaxel or with other PTP inducers and the release of 

cyt c in cell-free systems compared to intact cells has been explained by the existence of several 

drug targets inside the cell. Making only a subset of the drug available for mitochondria. 54 Likewise, 

other anti-tubulin agents such as vinorelbine or nocodazole could also been shown to trigger 

the release of cyt c via the direct interaction with mitochondria 101 that subsequently resulted in 

apoptotic cell death. The detection of tubulin as an inherent component of mitochondrial 

membranes able to interact with the voltage-dependent anion channel 102 and thereby able to 

influence the mPTPC seems to make the rather surprising identification of mitochondria as a 

target for well-established anti-tubulin agents plausible. 

For encapsulation of paclitaxel into DQAsomes, dequalinium chloride and paclitaxel were 

dissolved in methanol followed by removing the organic solvent. 103 After adding buffer, the 

suspension was sonicated with a probe sonicator until a clear opaque solution was formed. To 

remove any undissolved material, the sample was centrifuged for 10 min at 3000 rpm. The solubility 

of paclitaxel in water at 258C at pH 7.4 is with 0.172 mg/L (0.2 mM) extremely low, making 

any separation procedure of non-encapsulated paclitaxel unnecessary. However, for control, a 

paclitaxel suspension was probe sonicated under identical conditions used for the incorporation 

of paclitaxel into DQAsomes but in the complete absence of dequalinium. As expected, upon 

centrifugation, no paclitaxel was detectable in the supernatant using UV spectroscopy at 

230 nm. 103 Following this procedure, paclitaxel can be incorporated into DQAsomes at a molar 

ratio paclitaxel to dequalinium of about 0.6. In comparison to the free drug, encapsulation of 

paclitaxel into DQAsomes increases the drug’s solubility by a factor of about 3000. 

Weight distribution analysis 

90% 

50% 

10% 

Dust 

0.00% 

200 600 1000 

Size [nm] 

FIGURE 38.4 Paclitaxel encapsulated into DQAsomes. Left panel: transmission electron microscopic image 

(uranyl acetate staining); middle panel: Size distribution; right panel: Cryo-electron microscopic image. (From 

Cheng, S. M., Pabba, S., Torchilin, V. P., Fowle, W., Kimpfler, A., Schubert, R. and Weissig, V. J. Drug Deliv. 

Sci. Technol., 15(1), 81–86, 2005. With permission.) 


794 

Paclitaxel / DQA (molar) 

0.9 

0.8 

0.7 

0.6 

0.5 

0.4 

0.3 

0.2 

0.1 

0 

Day 1 Day 3 

Day 4 Day 6 Day 10 


FIGURE 38.5 Paclitaxel remains stably incorporated in DQAsomes: Molar ratio of paclitaxel to DQA in 

DQAsomes upon storage at 4C. (From Cheng, S. M., Pabba, S., Torchilin, V. P., Fowle, W. Kimpfler, A. 

Schubert, R., and Weissig, V., J. Drug Deliv. Sci. Technol., 15(1), 81–86, 2005. With permission.) 

Considering the known spherical character of DQAsomes, the results of an EM analysis of 

paclitaxel-loaded DQAsomes seem rather surprising. The transmission EM image (Figure 38.4, left 

panel) and the cryo-EM image (Figure 38.4, right panel) of an identical sample show with a 

remarkable conformity worm- or rod-like structures roughly around 400 nm in length, the size 

of which could also be confirmed by size distribution analysis shown in Figure 38.4 (middle panel). 

These complexes may represent the formation of worm-like micelles as recently described for selfassembling 

amphiphilic block co-polymers. 104 Figure 38.5 shows data about the stability of 

paclitaxel-containing DQAsomes upon storage. Over a period of ten days, the ratio of paclitaxel 

to dequalinium in the DQAsome preparation remains almost constant, indicating that paclitaxel 

remains stably incorporated in DQAsomes. However, from a gradual decrease of the total concentration 

of paclitaxel and of dequalinium over several days and to the same extent (by approximately 

20–30% after four days, data not shown) can be concluded that paclitaxel-loaded DQAsomes tend 

to form larger aggregates that are removable by centrifugation. This tendency to form aggregates is 

well known from artificial phospholipids vesicles (liposomes) and is generally reversible by 

shaking or slightly vortexing the preparation. 

The stability of paclitaxel-loaded DQAsomes is under physiological conditions in comparison 

to their shelf-life stability significantly reduced. In general, upon transferring paclitaxel-loaded 

DQAsomes either into serum-free cell medium (DMEM) at 378C or into complete serum, they 

form large aggregates leading to precipitation. However, this process of aggregation seems to be 

kinetically controlled, giving at least a portion of monomeric paclitaxel-loaded DQAsomes sufficient 

time to reach their target. The tendency to form aggregates under physiological conditions is 

common to all cationic carrier systems. Nevertheless, it should be mentioned that cationic lipoplexes 

and polyplexes are being tested in numerous gene therapeutic clinical trials. Also, cationic 

colloidal carriers (Catioms) are currently being successfully developed for systemic drug administrations. 

105–107 

38.7 DQASOMAL-ENCAPSULATED PACLITAXEL TRIGGERS APOPTOSIS 

IN VITRO 

To test if DQAsomal-encapsulated paclitaxel triggers apoptosis at paclitaxel concentrations where 

the free drug does not have a significant cytotoxic effect, human colo 205 colon cancer cells were 

incubated with the free drug, with empty DQAsomes, with a mixture of empty DQAsomes and 

the free drug, and with the DQAsomal-encapsulated drug. Following the staining of the treated cells 



% of apoptosis cells 

25 

20 

15 

10 

5 

0 

Empty DQAsomes mixture DQAsomes + 

free paclitaxel 

DQAsomal 

encapsulated 

paclitaxel 

FIGURE 38.6 Human colo 25 colon cancer cells were treated in triplicates with empty DQAsomes (20 nM 

DQA) with a mixture of empty DQAsomes (20 nM) and free paclitaxel (10 nM) and DQAsomal-encapsulated 

paclitaxel (20 nM DQA/10 nM paclitaxel). In each case, approximately 400 cells were counted (From Chen, S. 

M. and Weissig, V., 2006, manuscript in preparation). 

with the DNA-binding fluorophore Hoechst 33258, apoptotic nuclei showing the typical apoptotic 

condensation and fragmentation of chromatin were counted and expressed as percent of the total 

number of nuclei. Figure 38.6 shows that under identical incubation conditions, 10 nM paclitaxelencapsulated 

in DQAsomes more than doubles the number of apoptotic nuclei in comparison to the 

control whee cells were treated with a mixture of empty DQAsomes and 10 nM free paclitaxel. 

Likewise, a DNA ladder caused by DNA fragmentation typical for apoptosis could be detected 

upon incubation of colon cancer cells with 10 nM DQAsomal-encapsulated paclitaxel but not upon 

incubation with the free drug either alone or in mixture with empty DQAsomes (Figure 38.7). 

Incubating the cells for the same period of time, the amount of free paclitaxel had to be increased at 

least 5-fold over the amount of DQAsomal-encapsulated drug in order to generate a DNA ladder 

(not shown). 

FIGURE 38.7 Human colo 25 colon cancer cells were incubated for 30 h with buffer (lane 2), 20 nM empty 

DQAsomes (lane 3), 10 nM free paclitaxel (lane 4), a mixture of 20 nM empty DQAsomes and 10 nM free 

paclitaxel (lane 5), and 20 nM DQAsomes with 10 nM encapsulated paclitaxel (lane 6). White arrow heads 

indicate the apoptotic DNA ladder (Chen, S. M. and Weissig, V., 2006, manuscript in preparation). 


796 

Considering that paclitaxel, generally known as an anti-microtubule agent, has recently been 

demonstrated to trigger apoptosis by directly acting on mitochondria, 54,55,58 it can be concluded 

from the data shown in Figure 38.6 and Figure 38.7 that encapsulating paclitaxel into a 

mitochondria-specific drug delivery system appears to increase the sub-cellular, i.e., mitochondrial, 

bioavailability of the drug. 

38.8 TUMOR GROWTH INHIBITION STUDY WITH PACLITAXEL-LOADED 

DQASOMES IN VIVO 

Paclitaxel-loaded DQAsomes were tested for their ability to inhibit the growth of human colon 

cancer cells in nude mice. 103 COLO-205 cells were inoculated s.c. into the left flank of nude mice 

that all formed palpable tumors within seven days after cell injection. For controls with free 

paclitaxel, the drug was suspended in 100% DMSO at 20 mM, stored at 48C, and immediately 

before use, diluted in warm medium. In all controls, the dose of free paclitaxel and empty 

DQAsomes, respectively, was adjusted according to the dose of paclitaxel and dequalinium 

given in the paclitaxel-loaded DQAsome sample. Because of the lack of any inhibitory effect on 

tumor growth, the dose was tripled after 1.5 weeks. Figure 38.8 shows that at concentrations where 

free paclitaxel and empty DQAsomes do not show any impact on tumor growth, paclitaxel-loaded 

DQAsomes (with paclitaxel and dequalinium concentrations identical to controls) seem to inhibit 

the tumor growth by about 50%. Correspondingly, the average tumor weight in the treatment 

group after sacrificing the animals after 26 days is approximately half of that in all controls 

(Figure 38.9). Although this result seems to suggest that DQAsomes might be able to increase 

the therapeutic potential of paclitaxel, the preliminary character of this first in vivo study should 

be emphasized. 

Average tumor volume [mm 3 ] 

1400 

1200 

1000 

800 

600 

400 

200 

0 

0 10 20 

Day after tumor implantation 

30 


Hepes buffer 

Free paclitaxel 

Empty DQAsomes 

Paclitaxel-loaded 

DQAsomes 

FIGURE 38.8 Tumor growth inhibition study in nude mice implanted with human colon cancer cells. The 

mean tumor volume from each group was blotted against the number of days. Each group involved eight 

animals. For clarity, error bars were omitted. Note that after 1.5 weeks, the dose, normalized for paclitaxel, was 

tripled in all treatment groups. (From Cheng, S. M., Pabba, S., Torchilin, V. P., Fowle, W., Kimpfler, A., 

Schubert, R., and Weissig, V., J. Drug Deliv. Sci. Technol., 15(1), 81–86, 2005. With permission.) 




DQAsomes and DQAsome-like vesicles have been established as the first mitochondria-specific 

cationic drug delivery system potentially able to deliver cytotoxic drugs selectively to mitochondria 

in cancer cells. However, the encapsulation of antineoplastic drugs into cationic vesicles has 

already been suggested in 1998. 108 It was found that upon intravenous injection of gold-labeled 

cationic liposomes, 32% associated with tumor endothelial cells, 53% were internalized into 

endosomes, and 15% were extravascular 20 min after injection. 108 Following these early experimental 

data, Munich Biotech AG (Neuried, Germany) is currently developing so-called Catioms for 

the systemic delivery of anti-cancer drugs to solid tumors. 105–107 However, whereas Munich 

Biotech is focusing on the targeting of the tumor vasculature, the DQAsomal strategy described 

in this chapter is aimed at the delivery of pro-apoptotic compounds to and into tumor 

cell mitochondria. 

REFERENCES 

Average tumor weight [g] 

2.000 

1.600 

1.200 

0.800 

0.400 

0.000 

Hepes buffer Free paclitaxel Empty Paclitaxel-Loaded 

DQAsomes DQAsomes 

FIGURE 38.9 Average tumor weight (nZ8) at time of sacrifice of nude mice implanted with human colon 

cancer cells. For treatment group (paclitaxel-loaded DQAsomes) versus all three control groups combined 

PZ0.054 (parametric Student t-test). (From Cheng, S. M., Pabba, S., Torchilin, V. P., Fowle, W., Kimpfler, 

A., Schubert, R., and Weissig, V., J. Drug Deliv. Sci. Technol., 15(1), 81–86, 2005. With permission.) 

1. Costantini, P., Jacotot, E., Decaudin, D., and Kroemer, G., Mitochondrion as a novel target of 

anticancer chemotherapy, J. Natl. Cancer Inst., 92(13), 1042–1053, 2000. 

2. Galati, G., Teng, S., Moridani, M. Y., Chan, T. S., and O’Brien, P. J., Cancer chemoprevention and 

apoptosis mechanisms induced by dietary polyphenolics, Drug Metabol. Drug Interact., 17(1–4), 

311–349, 2000. 

3. Ghafourifar, P., Bringold, U., Klein, S. D., and Richter, C., Mitochondrial nitric oxide synthase, 

oxidative stress and apoptosis, Biol. Signals Recept., 10(1–2), 57–65, 2001. 

4. Davis, W., Ronai, Z., and Tew, K. D., Cellular thiols and reactive oxygen species in drug-induced 

apoptosis, J. Pharmacol. Exp. Ther., 296(1), 1–6, 2001. 

5. Creagh, E. M., Sheehan, D., and Cotter, T. G., Heat shock proteins—modulators of apoptosis in 

tumour cells, Leukemia, 14(7), 1161–1173, 2000. 

6. Hall, A. G., Review: The role of glutathione in the regulation of apoptosis, Eur. J. Clin. Invest., 

29(3), 238–245, 1999. 


798 


7. Deigner, H. P. and Kinscherf, R., Modulating apoptosis: Current applications and prospects for future 

drug development, Curr. Med. Chem., 6(5), 399–414, 1999. 

8. Peltenburg, L. T., Radiosensitivity of tumor cells. Oncogenes and apoptosis, Q. J. Nucl. Med., 44(4), 

355–364, 2000. 

9. Deigner, H. P., Haberkorn, U., and Kinscherf, R., Apoptosis modulators in the therapy of neurodegenerative 

diseases, Expert Opin. Investig. Drugs, 9(4), 747–764, 2000. 

10. Kanduc, D., Mittelman, A., Serpico, R., Sinigaglia, E., Sinha, A. A., Natale, C., Santacroce, R., 

Di Corcia, M. G. et al., Cell death: Apoptosis versus necrosis (review), Int. J. Oncol., 21(1), 

165–170, 2002. 

11. Robertson, J. D., Orrenius, S., and Zhivotovsky, B., Review: Nuclear events in apoptosis, J. Struct. 

Biol., 129(2,3), 346–358, 2000. 

12. Tomei, L. D. and Umansky, S. R., Apoptosis and the heart: A brief review, Ann. N. Y. Acad. Sci., 946, 

160–168, 2001. 

13. Sathasivam, S., Ince, P. G., and Shaw, P. J., Apoptosis in amyotrophic lateral sclerosis: A review of 

the evidence, Neuropathol. Appl. Neurobiol., 27(4), 257–274, 2001. 

14. Tinaztepe, K., Ozen, S., Gucer, S., and Ozdamar, S., Apoptosis in renal disease: A brief review of the 

literature and report of preliminary findings in childhood lupus nephritis, Turk. J. Pediatr., 43(2), 

133–138, 2001. 

15. Zusman, I., Gurevich, P., Gurevich, E., and Ben-Hur, H., The immune system, apoptosis and 

apoptosis-related proteins in human ovarian tumors (a review), Int. J. Oncol., 18(5), 965–972, 2001. 

16. Belzacq, A. S., Vieira, H. L., Kroemer, G., and Brenner, C., The adenine nucleotide translocator in 

apoptosis, Biochimie, 84(2,3), 167–176, 2002. 

17. Kaufmann, S. H. and Gores, G. J., Apoptosis in cancer: Cause and cure, Bioessays, 22(11), 

1007–1017, 2000. 

18. Olson, M. and Kornbluth, S., Mitochondria in apoptosis and human disease, Curr. Mol. Med., 1(1), 

91–122, 2001. 

19. Lin, Q. S., Mitochondria and apoptosis, Sheng Wu Hua Xue Yu Sheng Wu Wu Li Xue Bao (Shanghai), 

31(2), 116–118, 1999. 

20. Pollack, M. and Leeuwenburgh, C., Apoptosis and aging: Role of the mitochondria, J. Gerontol. A 

Biol. Sci. Med. Sci., 56(11), B475–B482, 2001. 

21. Henkart, P. A. and Grinstein, S., Apoptosis: Mitochondria resurrected? J. Exp. Med., 183(4), 

1293–1295, 1996. 

22. Skulachev, V. P., Why are mitochondria involved in apoptosis? Permeability transition pores and 

apoptosis as selective mechanisms to eliminate superoxide-producing mitochondria and cell, FEBS 

Lett., 397(1), 7–10, 1996. 

23. Petit, P. X., Zamzami, N., Vayssiere, J. L., Mignotte, B., Kroemer, G., and Castedo, M., Implication 

of mitochondria in apoptosis, Mol. Cell. Biochem., 174(1,2), 185–188, 1997. 

24. Decaudin, D., Marzo, I, Brenner, C., and Kroemer, G., Mitochondria in chemotherapy-induced 

apoptosis: A prospective novel target of cancer therapy (review), Int. J. Oncol., 12(1), 141–152, 

1998. 

25. Mignotte, B. and Vayssiere, J. L., Mitochondria and apoptosis, Eur. J. Biochem., 252(1), 1–15, 1998. 

26. Susin, S. A., Zamzami, N., and Kroemer, G., Mitochondria as regulators of apoptosis: Doubt no 

more, Biochim. Biophys. Acta, 1366(1,2), 151–165, 1998. 

27. Green, D. R. and Reed, J. C., Mitochondria and apoptosis, Science, 281(5381), 1309–1312, 1998. 

28. Tatton, W. G. and Chalmers-Redman, R. M., Mitochondria in neurodegenerative apoptosis: An 

opportunity for therapy? Ann. Neurol., 44(3 Suppl. 1), S134–S141, 1998. 

29. Gottlieb, R. A., Mitochondria: Ignition chamber for apoptosis, Mol. Genet. Metab., 68(2), 227–231, 

1999. 

30. Nieminen, A. L., Apoptosis and necrosis in health and disease: Role of mitochondria, Int. Rev. 

Cytol., 224, 29–55, 2003. 

31. Weber, T., Dalen, H., Andera, L., Negre-Salvayre, A., Auge, N., Sticha, M., Lloret, A. et al., 

Mitochondria play a central role in apoptosis induced by alpha-tocopheryl succinate, an agent 

with antineoplastic activity: Comparison with receptor-mediated pro-apoptotic signaling, Biochemistry, 

42(14), 4277–4291, 2003. 



32. Richter, C., Oxidative Stress, mitochondria, and apoptosis, Restor. Neurol. Neurosci., 12(2,3), 

59–62, 1998. 

33. Gulbins, E., Dreschers, S., and Bock, J., Role of mitochondria in apoptosis, Exp. Physiol., 88(Pt 1), 

85–90, 2003. 

34. Hockenbery, D. M., Giedt, C. D., O’Neill, J. W., Manion, M. K., and Banker, D. E., Mitochondria 

and apoptosis: New therapeutic targets, Adv. Cancer Res., 85, 203–242, 2002. 

35. Heiligtag, S. J., Bredehorst, R., and David, K. A., Key role of mitochondria in cerulenin-mediated 

apoptosis, Cell Death Differ., 9(9), 1017–1025, 2002. 

36. Boichuk, S. V., Minnebaev, M. M., and Mustafin, I. G., Key role of mitochondria in apoptosis of 

lymphocytes, Bull. Exp. Biol. Med., 132(6), 1166–1168, 2001. 

37. Birbes, H., Bawab, S. E., Obeid, L. M., and Hannun, Y. A., Mitochondria and ceramide: Intertwined 

roles in regulation of apoptosis, Adv. Enzyme Regul., 42, 113–129, 2002. 

38. Naoi, M., Maruyama, W., Akao, Y., and Yi, H., Mitochondria determine the survival and death in 

apoptosis by an endogenous neurotoxin, N-methyl(R)salsolinol, and neuroprotection by propargylamines, 

J. Neural Transm., 109(5,6), 607–621, 2002. 

39. Roucou, X., Antonsson, B., and Martinou, J. C., Involvement of mitochondria in apoptosis, Cardiol. 

Clin., 19(1), 45–55, 2001. 

40. Wang, X., The expanding role of mitochondria in apoptosis, Genes Dev., 15(22), 2922–2933, 

2001. 

41. Raha, S. and Robinson, B. H., Mitochondria, oxygen free radicals, and apoptosis, Am. J. Med. 

Genet., 106(1), 62–70, 2001. 

42. Waterhouse, N. J., Goldstein, J. C., Kluck, R. M., Newmeyer, D. D., and Green, D. R., The (Holey) 

study of mitochondria in apoptosis, Methods Cell Biol., 66, 365–391, 2001. 

43. Sordet, O., Rebe, C., Leroy, I., Bruey, J. M., Garrido, C., Miguet, C., Lizard, G., Plenchette, S., 

Corcos, L., and Solary, E., Mitochondria-targeting drugs arsenic trioxide and lonidamine bypass the 

resistance of TPA-differentiated leukemic cells to apoptosis, Blood, 97(12), 3931–3940, 2001. 

44. Gottlieb, R. A., Mitochondria and apoptosis, Biol. Signals Recept., 10(3,4), 147–161, 2001. 

45. Shi, Y., A structural view of mitochondria-mediated apoptosis, Nat. Struct. Biol., 8(5), 394–401, 

2001. 

46. Gottlieb, R. A., Role of mitochondria in apoptosis, Crit. Rev. Eukaryot. Gene. Expr., 10(3,4), 

231–239, 2000. 

47. Jiang, Z. F., Zhao, Y., Hong, X., and Zhai, Z. H., Nuclear apoptosis induced by isolated mitochondria, 

Cell Res., 10(3), 221–232, 2000. 

48. Brenner, C. and Kroemer, G., Apoptosis. Mitochondria—the death signal integrators, Science, 

289(5482), 1150–1151, 2000. 

49. Desagher, S. and Martinou, J. C., Mitochondria as the central control point of apoptosis, Trends Cell 

Biol., 10(9), 369–377, 2000. 

50. Kroemer, G., Mitochondrial control of apoptosis: An overview, In Mitochondria and Cell Death, 

Brown, G. C., Nicholls, D. G., and Cooper, C. E., Eds., Princton University Press, Princeton, NJ, 

pp. 1–15, 1999. 

51. Fulda, S., Susin, S. A., Kroemer, G., and Debatin, K. M., Molecular ordering of apoptosis induced by 

anticancer drugs in neuroblastoma cells, Cancer Res., 58(19), 4453–4460, 1998. 

52. Ferlini, C., Raspaglio, G., Mozzetti, S., Distefano, M., Filippetti, F., Martinelli, E., Ferrandina, G., 

Gallo, D., Ranelletti, F. O., and Scambia, G., Bcl-2 down-regulation is a novel mechanism of 

paclitaxel resistance, Mol. Pharmacol., 64(1), 51–58, 2003. 

53. von Haefen, C., Wieder, T., Essmann, F., Schulze-Osthoff, K., Dorken, B., and Daniel, P. T., 

Paclitaxel-induced apoptosis in BJAB cells proceeds via a death receptor-independent, caspases- 

3/-8-driven mitochondrial amplification loop, Oncogene, 22(15), 2236–2247, 2003. 

54. Andre, N., Carre, M., Brasseur, G., Pourroy, B., Kovacic, H., Briand, C., and Braguer, D., Paclitaxel 

targets mitochondria upstream of caspase activation in intact human neuroblastoma cells, FEBS 

Lett., 532(1–2), 256–260, 2002. 

55. Kidd, J. F., Pilkington, M. F., Schell, M. J., Fogarty, K. E., Skepper, J. N., Taylor, C. W., and Thorn, P., 

Paclitaxel affects cytosolic calcium signals by opening the mitochondrial permeability transition pore, 

J. Biol. Chem., 277(8), 6504–6510, 2002. 


800 


56. Makin, G. W., Corfe, B. M., Griffiths, G. J., Thistlethwaite, A., Hickman, J. A., and Dive, C., 

Damage-induced Bax N-terminal change, translocation to mitochondria and formation of Bax 

dimers/complexes occur regardless of cell fate, Embo J., 20(22), 6306–6315, 2001. 

57. Varbiro, G., Veres, B., Gallyas, F., and Sumegi, B., Direct effect of Taxol on free radical formation 

and mitochondrial permeability transition, Free Radic. Biol. Med., 31(4), 548–558, 2001. 

58. Andre, N., Braguer, D., Brasseur, G., Goncalves, A., Lemesle-Meunier, D., Guise, S., Jordan, M. A., 

and Briand, C., Paclitaxel induces release of cytochrome c from mitochondria isolated from human 

neuroblastoma cells’, Cancer Res., 60(19), 5349–5353, 2000. 

59. Carre, M., Carles, G., Andre, N., Douillard, S., Ciccolini, J., Briand, C., and Braguer, D., Involvement 

of microtubules and mitochondria in the antagonism of arsenic trioxide on paclitaxel-induced 

apoptosis, Biochem. Pharmacol., 63(10), 1831–1842, 2002. 

60. Perkins, C. L., Fang, G., Kim, C. N., and Bhalla, K. N., The role of Apaf-1, caspase-9, and bid 

proteins in etoposide- or paclitaxel-induced mitochondrial events during apoptosis, Cancer Res., 

60(6), 1645–1653, 2000. 

61. Itoh, M., Noutomi, T., Toyota, H., and Mizuguchi, J., Etoposide-mediated sensitization of squamous 

cell carcinoma cells to tumor necrosis factor-related apoptosis-inducing ligand (TRAIL)-induced 

loss in mitochondrial membrane potential, Oral Oncol., 39(3), 269–276, 2003. 

62. Custodio, J. B., Cardoso, C. M., Madeira, V. M., and Almeida, L. M., Mitochondrial 

permeability transition induced by the anticancer drug etoposide, Toxicol. In Vitro, 15(4,5), 

265–270, 2001. 

63. Fujino, M., Li, X. K., Kitazawa, Y., Guo, L., Kawasaki, M., Funeshima, N., Amano, T., and Suzuki, S., 

Distinct pathways of apoptosis triggered by FTY720, etoposide, and anti-Fas antibody in human 

T-lymphoma cell line (Jurkat cells), J. Pharmacol. Exp. Ther., 300(3), 939–945, 2002. 

64. Robertson, J. D., Gogvadze, V., Zhivotovsky, B., and Orrenius, S., Distinct pathways for stimulation 

of cytochrome c release by etoposide, J. Biol. Chem., 275(42), 32438–32443, 2000. 

65. Marchetti, P., Zamzami, N., Joseph, B., Schraen-Maschke, S., Mereau-Richard, C., Costantini, P., 

Metivier, D., Susin, S. A., Kroemer, G., and Formstecher, P., The novel retinoid 6-[3-(1-adamantyl)- 

4-hydroxyphenyl]-2-naphtalene carboxylic acid can trigger apoptosis through a mitochondrial 

pathway independent of the nucleus, Cancer Res., 59(24), 6257–6266, 1999. 

66. Murphy, M. P., Slip and leak in mitochondrial oxidative phosphorylation, Biochim. Biophys. Acta, 

977, 123–141, 1989. 

67. Murphy, M. P. and Smith, R. A., Drug delivery to mitochondria: The key to mitochondrial medicine, 


68. Modica-Napolitano, J. S. and Aprille, J. R., Delocalized lipophilic cations selectively target the 

mitochondria of carcinoma cells, Adv. Drug Deliv. Rev., 49(1,2), 63–70, 2001. 

69. Modica-Napolitano, J. S. and Aprille, J. R., Basis for the selective cytotoxicity of rhodamine 123, 

Cancer Res., 47(16), 4361–4365, 1987. 

70. Modica-Napolitano, J. S. and Singh, K. K., Mitochondria as targets for detection and treatment of 

cancer,Expert Rev. Mol. Med. 2002 April 11; 2002:1–19. 

71. Modica-Napolitano, J. S., Koya, K., Weisberg, E., Brunelli, B. T., Li, Y., and Chen, L. B., Selective 

damage to carcinoma mitochondria by the rhodacyanine MKT-077, Cancer Res., 56(3), 544–550, 

1996. 

72. Manetta, A., Emma, D., Gamboa, G., Liao, S., Berman, M., and DiSaia, P., Failure to enhance the 

in vivo killing of human ovarian carcinoma by sequential treatment with dequalinium chloride and 

tumor necrosis factor, Gynecol. Oncol., 50(1), 38–44, 1993. 

73. Weiss, M. J., Wong, J. R., Ha, C. S., Bleday, R., Salem, R. R., Steele, G. D., and Chen, L. B., 

Dequalinium, a topical antimicrobial agent, displays anticarcinoma activity based on selective 

mitochondrial accumulation, Proc. Natl. Acad. Sci. USA, 84(15), 5444–5448, 1987. 

74. Christman, J. E., Miller, D. S., Coward, P., Smith, L. H., and Teng, N. N., Study of the selective 

cytotoxic properties of cationic, lipophilic mitochondrial-specific compounds in gynecologic malignancies, 

Gynecol. Oncol., 39(1), 72–79, 1990. 

75. Davis, S., Weiss, M. J., Wong, J. R., Lampidis, T. J., and Chen, L. B., Mitochondrial and plasma 

membrane potentials cause unusual accumulation and retention of rhodamine 123 by human breast 

adenocarcinoma-derived MCF-7 cells, J. Biol. Chem., 260(25), 13844–13850, 1985. 



76. Koya, K., Li, Y., Wang, H., Ukai, T., Tatsuta, N., Kawakami, M., Shishido, T., and Chen, L. B., 

MKT-077, a novel rhodacyanine dye in clinical trials, exhibits anticarcinoma activity in preclinical 

studies based on selective mitochondrial accumulation, Cancer Res., 56(3), 538–543, 1996. 

77. Weisberg, E. L., Koya, K., Modica-Napolitano, J., Li, Y., and Chen, L. B., In vivo administration of 

MKT-077 causes partial yet reversible impairment of mitochondrial function, Cancer Res., 56(3), 

551–555, 1996. 

78. Chiba, Y., Kubota, T., Watanabe, M., Matsuzaki, S. W., Otani, Y., Teramoto, T., Matsumoto, Y., 

Koya, K., and Kitajima, M., MKT-077, localized lipophilic cation: Antitumor activity against human 

tumor xenografts serially transplanted into nude mice, Anticancer Res., 18(2A), 1047–1052, 1998. 

79. Chiba, Y., Kubota, T., Watanabe, M., Otani, Y., Teramoto, T., Matsumoto, Y., Koya, K., and 

Kitajima, M., Selective antitumor activity of MKT-077, a delocalized lipophilic cation, on normal 

cells and cancer cells in vitro, J. Surg. Oncol., 69(2), 105–110, 1998. 

80. Petit, T., Izbicka, E., Lawrence, R. A., Nalin, C., Weitman, S. D., and Von Hoff, D. D., Activity of 

MKT 077, a rhodacyanine dye, against human tumor colony-forming units, Anticancer Drugs, 10(3), 

309–315, 1999. 

81. Propper, D. J., Braybrooke, J. P., Taylor, D. J., Lodi, R., Styles, P., Cramer, J. A., Collins, W. C. 

et al., Phase I trial of the selective mitochondrial toxin MKT077 in chemo-resistant solid tumours, 

Ann. Oncol., 10(8), 923–927, 1999. 

82. Adams, D. M., Ji, W., Barth, R. F., and Tjarks, W., Comparative in vitro evaluation of dequalinium 

B, a new boron carrier for neutron capture therapy (NCT), Anticancer Res., 20(5B), 3395–3402, 

2000. 

83. Weissig, V., Mogel, H. J., Wahab, M., and Lasch, J., Computer simulations of DQAsomes, Proceed. 

Intl. Symp. Control. Rel. Bioact. Mater., 25, 312, 1998. 

84. Weissig, V., Lasch, L., Erdos, G., Meyer, H. W., Rowe, T. C., and Hughes, J., DQAsomes: A novel 

potential drug and gene delivery system made from Dequalinium, Pharm. Res., 15(2), 334–337, 

1998. 

85. De Rosa, M., Gambacorta, A., and Gliozi, A., Structure, biosynthesis, and physicochemical properties 

of archaebacterial lipds, Microbiol. Rev., 50, 70–80, 1986. 

86. Gambacorta, A., Gliozi, A., and De Rosa, M., Archaeal lipids and their biotechnological applications, 

World J. Microbiol. Biotechnol., 11, 115–131, 1995. 

87. Weissig, V. and Torchilin, V. P., Mitochondriotropic cationic vesicles: A strategy towards mitochondrial 

gene therapy, Curr. Pharm. Biotechnol., 1(4), 325–346, 2000. 

88. Weissig, V. and Torchilin, V. P., Cationic single-chain bolaamphiphiles as new materials for application 

in medicine and biotechnology. In Materials research society meeting, Boston, MA., EE.5.4, 

p. 173, 1999. 

89. Weissig, V., Lizano, C., Ganellin, C. R., and Torchilin, V. P., DNA binding cationic bolasomes with 

delocalized charge center: A structure-activity relationship study, STP Pharma Sci., 11, 91–96, 2001. 

90. Lasch, J., Meye, A., Taubert, H., Koelsch, R., Mansa-ard, J., and Weissig, V., Dequalinium vesicles 

form stable complexes with plasmid DNA which are protected from DNase attack, Biol. Chem., 

380(6), 647–652, 1999. 

91. Weissig, V., Lizano, C., and Torchilin, V. P., Selective DNA release from DQAsome/DNA 

complexes at mitochondria-like membranes, Drug Deliv., 7(1), 1–5, 2000. 

92. Weissig, V., D’Souza, G. G., and Torchilin, V. P., DQAsome/DNA complexes release DNA upon 

contact with isolated mouse liver mitochondria, J. Control Release, 75(3), 401–408, 2001. 

93. Ravagnan, L., Marzo, I., Costantini, P., Susin, S. A., Zamzami, N., Petit, P. X., and Hirsch, F., 

Lonidamine triggers apoptosis via a direct, Bcl-2-inhibited effect on the mitochondrial permeability 

transition pore., Oncogene, 18(16), 2537–2546, 1999. 

94. Jacotot, E., Costantini, P., Laboureau, E., Zamzami, N., Susin, S. A., and Kroemer, G., Mitochondrial 

membrane permeabilization during the apoptotic process, Ann. N. Y. Acad. Sci., 887, 

18–30, 1999. 

95. van der Woude, I., Wagenaar, A., Meekel, A. A., ter Beest, M. B., Ruiters, M. H., Engberts, J. B., and 

Hoekstra, D., Novel pyridinium surfactants for efficient, nontoxic in vitro gene delivery, Proc. Natl. 

Acad. Sci. USA, 94(4), 1160–1165, 1997. 

96. Zuhorn, I. S. and Hoekstra, D., On the mechanism of cationic amphiphile-mediated transfection. 

To fuse or not to fuse: Is that the question? J. Membr. Biol., 189(3), 167–179, 2002. 


802 


97. Wattiaux, R., Jadot, M., Warnier-Pirotte, M. T., and Wattiaux-De Coninck, S., Cationic lipids 

destabilize lysosomal membrane in vitro, FEBS Lett., 417(2), 199–202, 1997. 

98. El Ouahabi, A., Thiry, M., Pector, V., Fuks, R., Ruysschaert, J. M., and Vandenbranden, M., The role 

of endosome destabilizing activity in the gene transfer process mediated by cationic lipids, FEBS 

Lett., 414(2), 187–192, 1997. 

99. D’Souza, G. G., Rammohan, R., Cheng, S. M., Torchilin, V. P., and Weissig, V., DQAsomemediated 

delivery of plasmid DNA toward mitochondria in living cells, J. Control Release, 92, 

189–197, 2003. 

100. D’Souza, G. G., Boddapati, S. V., and Weissig, V., Mitochondrial leader sequence-plasmid DNA 

conjugates delivered into mammalian cells by DQAsomes co-localize with mitochondria, Mitochondrion, 

5(5), 352–358, 2005. 

101. Braguer, D., Andre, N., Carre, M., Carles, G., Gonzalves, A., and Briand, C., 92nd Annual Meeting of 

the American Association for Cancer Research, American Association for Cancer Research, New 

Orleans, LA, p. 369, 2001. 

102. Carre, M., Andre, N., Carles, G., Borghi, H., Brichese, L., Briand, C., and Braguer, D., Tubulin is an 

inherent component of mitochondrial membranes that interacts with the voltage-dependent anion 

channel, J. Biol. Chem., 277, 33664–33669, 2002. 

103. Cheng, S. M., Pabba, S., Torchilin, V. P., Fowle, W., Kimpfler, A., Schubert, R., and Weissig, V., 

Towards mitochondria-specific delivery of apoptosis-inducing agents: DQAsomal incorporated 

paclitaxel, J. Drug Deliv. Sci. Techno., 15(1), 81–86, 2005. 

104. Discher, E. D. and Eisenberg, A., Polymer vesicles, Science, 297, 967–973, 2002. 

105. Schuppenhauer, M., Munich Biotech: Catioms for cancer. BioCentury, The Bernstein Report on 

BioBusiness, May 19, A10, 2003. 

106. Schmitt-Sody, M., Strieth, S., Krasnici, S., Sauer, B., Schulze, B., Teifel, M., Michaelis, U., Naujoks, K., 

and Dellian, M., Neovascular targeting therapy: Paclitaxel encapsulated in cationic liposomes improves 

antitumoral efficacy, Clinical Cancer Research, 9(6), 2335–2341, 2003. 

107. Krasnici, S., Werner, A., Eichhorn, M. E., Schmitt-Sody, M., Pahernik, S. A., Sauer, B., and Schulze, B., 

Effect of the surface charge of liposomes on their uptake by angiogenic tumor vessels, Int. J. Cancer, 

105(4), 561–567, 2003. 

108. Thurston, G., McLean, J. W., Rizen, M., Baluk, P., Haskell, A., Murphy, T. J., Hanahan, D., and 

McDonald, D. M., Cationic liposomes target angiogenic endothelial cells in tumors and chronic 

inflammation in mice, J. Clin. Invest., 101(7), 1401–1413, 1998. 


Section 1 

Nanotechnology and Cancer 


1

Section 2 

Polymer Conjugates 


157

Section 3 

Polymeric Nanoparticles 


213

Section 4 

Polymeric Micelles 


315

Section 5 

Dendritic Nanocarriers 


487

Section 6 

Liposomes 


593

Section 7 

Other Lipid Nanostructures 


721

q 2006 by Taylor & Francis Group, LLC - Developers

You also want an ePaper? Increase the reach of your titles

Delete template?

Save as template?